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UNIT III Syllabus: Polymeric implant materials: Polyolefins, polyamides, acrylic polymers, fluorocarbon polymers. Rubbers, acetals.(Classification according to thermosets, Thermoplastics and elastomers) . Physiochemical characteristics of biopolymers. Biodegradable polymers for medical purposes. Biopolymers in controlled release systems. Synthetic polymeric membranes and their biological applications. Polymeric implant materials:

Synthetic polymeric materials have been widely used in medical disposable supply, prosthetic materials, dental materials, implants, dressings, extracorporeal devices, encapsulants, polymeric drug delivery systems, tissue engineered products, and orthodoses as that of metal and ceramics substituents.

The main advantages of the polymeric biomaterials compared to metal or ceramic materials are ease of manufacturability to produce various shapes (latex, film, sheet, fibers, etc.), ease of secondary processability, reasonable cost, and availability with desired mechanical and physical properties.

The required properties of polymeric biomaterials are similar to other biomaterials, that is, biocompatibility, sterilizability, adequate mechanical and physical properties, and manufacturability

Polymerization and Basic Structure: Polymerization

In order to link the small molecules one has to force them to lose their electrons by the chemical processes of condensation and addition. By controlling the reaction temperature, pressure, and time in the presence of catalyst(s), the degree to which repeating units are put together into chains can be manipulated.

Requirements for Biomedical Polymers . Properties Description .

Compatibility : Noncarcinogenesis, nonpyrogenicity, nontoxicity, and nonallergic response Sterilizability : Autoclave, dry heating, ethylene oxide gas, and radiation Physical property : Strength, elasticity, and durability Manufacturability : Machining, molding, extruding, and fiber forming . Typical Condensation Polymers Type Inter-unit Linkage Polyester O

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–C–O– Polyamide O H

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–C–N– Polyurea H O H

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–N–C–N– Polyurethane O H

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–O–C–N–

Polysiloxane R |

–Si–O– 1

R

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Protein O H || |

–C–N–

Polyolefin

Polyolefins is the collective name for polyethylene (PE) and polypropylene (PP) - among the most widely used plastics today. They account for more than half of total plastics consumption in the developed world. Polyolefins are extremely versatile. They posess a successful combination of properties, including

• flexibility • strength • lightness • stability • impermeability, and • easy processability.

PE and PP are also well-suited to recycling and re-use. Cracking The production chain begins with oil and gas. "Cracking" refinery-derived petroleum products at high temperature produces ethylene and propylene, collectively called olefins. These gases are the building blocks for a large number of chemical products including polyolefin plastics.

Polymerisation The next stage, polymerisation, converts ethylene and propylene molecules into long chains known as polymers. Thus, ethylene is polymerised into polyethylene and propylene into polypropylene. Modifying this process yields an extremely broad range of PE and PP products with differing properties. A polyolefin is a polymer produced from a simple olefin, or alkene as a monomer. For example, polyethylene is the polyolefin produced by polymerizing the olefin ethylene. An equivalent term is polyalkene; this is a more modern term, although polyolefin is still used in the petrochemical industry. Polypropylene is another common polyolefin which is made from propylene. A more specific type of olefin is a poly-alpha-olefin (or poly-α-olefin, sometimes abbreviated as PAO), a polymer made by polymerizing an alpha-olefin. An alpha-olefin (or α-olefin) is an alkene where the carbon-carbon double bond starts at the α-carbon atom, i.e. the double bond is between the #1 and #2 carbons in the molecule. Common alpha-olefins used as co-monomers to give a polymer alkyl branching groups are similar to 1-hexene or may be longer (see chemical structure below).

1-hexene, an example of an alpha-olefin Many poly-alpha-olefins have flexible alkyl branching groups on every other carbon of their polymer backbone chain. These alkyl groups, which can shape themselves in numerous conformations, make it very difficult for the polymer molecules to line themselves up side-by-side in an orderly way. Therefore, many poly-alpha-olefins do not crystallize or solidify easily and are able to remain oily, viscous liquids even at lower temperatures. Low molecular weight poly-alpha-olefins are useful as synthetic lubricants such as synthetic motor oils for vehicles used in a wide temperature range. Even polyethylenes copolymerized with a small amount of alpha-olefins (such as 1-hexene, 1-octene, or longer) are more flexible than simple straight chain high density polyethylene, which

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has no branching. The methyl branch groups on a polypropylene polymer are not long enough to make typical commercial polypropylene more flexible than polyethylene. Polyamide A polyamide is a polymer containing monomers joined by peptide bonds. They can occur both naturally, examples being proteins, such as wool and silk, and can be made artificially, examples being Nylons, Aramids, and sodium poly(aspartate). Production from monomers The amide link is produced from the condensation reaction of an amino group and a carboxylic acid or acid chloride group. A small molecule, usually water, ammonia or hydrogen chloride, is eliminated. The amino group and the carboxylic acid group can be on the same monomer, or the polymer can be constituted of two different bifunctional monomers, one with two amino groups, the other with two carboxylic acid or acid chloride groups. Amino Acids can be taken as examples of single monomer (if the difference between R groups is ignored) reacting with identical molecules to form a polyamide:

The reaction of two amino acids. Many of these reactions produce long chain proteins Aramid (pictured below) is made from two different monomers which continuously alternate to form the polymer and is an aromatic polyamide:

The reaction of 1,4-phenyl-diamine (para-phenylenediamine) and terephthaloyl chloride to produce Aramid Nylon Acrylic polymer:

Acrylic is a generic term denoting derivatives of acrylic and methacrylic acid, including acrylic esters and compounds containing nitrile and amide groups. Polymers based on acrylics were discovered before many other polymers that are now widely employed. I

Optical clarity is an important quality in medical devices or diagnostic equipment that rely on visual inspection and therefore require a high level of transparency.

Acrylic, the common name for polymethyl methacrylate (PMMA), has long been used in the manufacture of cuvettes, tubing connectors, speculums, and many other medical devices requiring impact strength, chemical resistance, biocompatibility, and clarity.

Acrylic polymers thus occupy a prominent place in the market for clear, disposable plastics--only glass transmits light as well. Acrylic also is commonly used in the manufacture of reading glasses, due to its superior optical qualities.

The properties of acrylic materials and compares them with other thermoplastics competing for a share of the medical market.

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Benefits to both the manufacturer and the end-user are highlighted, as are applications for which acrylic is well suited and those for which it is not recommended.

In addition, the paper discusses acrylic's chemical resistance and performance after sterilization, along with disposal, recycling, and other environmental issues. Design and processing guidelines for manufacturers are also included as is a case study of a specific application. Overview of Acrylic

Acrylic polymer, derived from the monomer methyl methacrylate (MMA), was first developed more than 60 years ago. General-purpose acrylic grades contain a co-monomer, added during the polymerization process, to facilitate flow during injection molding and extrusion. Specialty grades are formulated to perform in applications requiring high impact strength and heat resistance.

UV-light-transmitting formulations are also available, and are specified for certain critical diagnostic equipment in which even slight UV absorption or variation in material flow could be detrimental.

Easily processed and assembled, acrylic has been used in medical and health-care applications since its introduction. One of the first uses of acrylic sheet was for incubators.

The first intraocular acrylic prosthesis was implanted in 1955, and ever since acrylic has been used in contact with human tissue.

Its biocompatibility led to the adoption of acrylic for aircraft canopies during World War II: pilots suffered fewer infections from shards of acrylic than they had from glass.

The leading applications of acrylic in the medical industry today are for cuvettes and tubing connectors, but it is also used to produce test kits, syringes, luers, blood filters, and drainage wands as well as flowmeters, blood-pump housings, fluid silos, surgical-blade dispensers, incubators, and surgical trays. Acrylic polymers are resistant to many biological and chemical agents.

The use of acrylic polymers in the medical industry has been steadily increasing over the past several years.

This is especially noticeable in the area of diagnostics, due to the fact that acrylic is an inert material that does not react with the reagents used in testing.

For medical devices, special impact-modified grades formulated to resist breaking and cracking are employed more often than standard grades.

Physical Properties

Acrylic offers light transmittance of 92%--theoretically the maximum obtainable--with particular clarity at lower wavelengths of 270 to 350 nm.

For example, acrylic is the material of choice for disposable cuvettes, used to contain blood and other fluids, through which a spectrum of UV light is passed for analysis.

Although quartz glass can be used for the most demanding applications, since it transmits light as low as 220 nm, it is not cost-effective in an application that increasingly favors disposable plastics.

Other prominent physical properties of acrylic include good mechanical strength and dimensional stability; along with high tensile and flexural strength .

Medical-grade acrylic can be supplied for intricate, thin-wall applications in which maximum optical clarity is required: it offers excellent injection molding flow properties (13 g/10 min) and a tensile strength of 10,600 psi.

Acrylic also provides good surface hardness for scratch resistance, an important quality in medical applications.

Because acrylic is a rigid material, standard grades do not provide high impact resistance. Therefore, impact-modified grades--softer and less rigid than standard formulations--are specified for applications that typically require increased toughness.

Acrylic is not recommended for applications that demand very high impact resistance or those that put surfaces under high pressure.

Acrylic does perform well in electrical applications, due to its insulating nature; an increase in absorbed moisture makes it more conductive.

As temperatures increase, acrylic becomes more flexible and exhibits less flexural strength. Under sustained loading, strain on the material can induce excessive molecular movement that

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increases with time under load and higher temperatures and results in the phenomenon known as creep that is common to all thermoplastics.

Chemical Resistance

Acrylic is resistant to a wide range of chemicals including salts, bases, aliphatic hydrocarbons, fats and oils, most common gases and inorganic chemicals, dilute mineral and organic acids, and dilute and concentrated solutions of most alkalis.

It is attacked by strong acids, chlorinated and aromatic hydrocarbons, ketones, alcohols, ethers, and esters. Sterilization Effects Acceptable sterilization technologies for acrylic are E-beam or gamma irradiation or dry ethylene oxide gas. While gamma sterilization has a tendency to discolor (yellow) most acrylics, this yellowing is temporary and recovery can be complete, with the parts retaining their original integrity. The higher the radiation dosage, the greater the yellowing and the longer the required recovery time. Material Selection

The failure of plastic parts can often be traced to mistakes in design, production methods, or material choice. Selecting the material for a particular application involves careful consideration of the end-use environmental conditions and functional requirements of the part.

Basing the selection on incorrect criteria can lead to overdesigning the part or specifying a material with properties that exceed the demands of the application, often at the expense of scratch or chemical resistance, clarity, or other desirable qualities.

Environmental conditions that figure into the selection include chemicals likely to be encountered, sterilization methods, humidity, temperature, and thermal cycling. Other important factors are mold and part design, processing requirements, and assembly methods.

In selecting the best grade of acrylic to use, these same factors must be considered. Higher-molecular-weight grades are more resistant to crazing from chemical exposure and mechanical stress, but have lower melt-flow rates.

As the level of impact modifier is increased in impact-grade acrylic, other properties such as clarity, light transmission, and tensile strength often diminish.

Processing requirements can be difficult to predict. Flow patterns and mold-filling problems can result from large part surfaces, deep ribs or flanges, asymmetrical geometries, and unbalanced thick and thin sections.

Experimenting with several different grades in the prototype tool is recommended. Molds designed for acrylic, ABS, and polycarbonate parts typically accommodate a shrinkage level of 0.006 in./in., allowing for convenient trial runs of acrylic in molds used for ABS or polycarbonate parts.

Direct comparison of the parts molded from different materials can help determine the best one for the application.

Processing Considerations

Compared with other polymers, acrylic is relatively easy to process. It can be molded with little or no residual stress and is available in formulations specifically designed for injection molding or extrusion in a wide range of melt-flow rates.

Under normal processing conditions, acrylic produces melts that are typically higher in viscosity than those of many other thermoplastic polymers. The higher-molecular-weight grades are generally recommended for extrusion.

Because of acrylic's higher- viscosity flow properties, injection molding runners and sprues used to process it need to have larger diameters than those handling polystyrene or polyethylene.

Large injection-molded parts or parts with thin-wall geometries may require a high melt-flow rate.

When acrylic is processed in molds built for polycarbonate or polyester, lower injection pressures should be used, as it is less likely to warp than polyesters and is more forgiving in molds not designed with well-balanced gates.

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Because acrylic has a higher melt viscosity than polystyrene or styrenic alloys, it requires higher injection pressures when selected in place of those materials.

Higher clamping forces may also be required. A hygroscopic material, acrylic absorbs water and must be dried prior to molding: if molded while wet, it exhibits moisture splay, leaving streaks, bubbles, and a rough surface on the part.

Typical wall thicknesses for acrylic parts range from 0.040 to 0.500 in. Thicker or thinner parts can be achieved with special designs or processing methods such as injection/ compression molding. Consistency is the key; any changes in thickness should be gradual, and feature radiused edges.

Vertical walls should be the same thickness as the rest of the part to avoid pressure variations on the flow front, which can lead to stressed areas and voids caused by trapped air.

Moderate residual stress does not affect part performance, but high levels of stress can reduce impact strength and resistance to chemical or heat crazing, and can undermine the dimensional stability of the molded part.

Before the part and mold design are completed, a mold-flow analysis should be performed to help avoid costly mistakes and downtime.

Bonding Systems

Acrylic parts can be fastened by chemical bonding, ultrasonic welding, and heat staking. Two types of agents are commonly used to chemically bond acrylic: solvents and polymerizable adhesives.

Solvents such as dichloromethane dissolve the surfaces of two acrylic parts, which harden after the solvent evaporates and bond to one another.

Solutions of acrylic polymer dissolved in a solvent or methyl-methacrylate monomer work similarly.

Two-part polymerizable adhesives contain a viscous acrylic resin base and a liquid catalyst that when mixed together provide a strong joint.

Ultrasonic welding is an efficient method of fusing two parts made from the same material. Both contact (near-field) welding and transmission (far-field) welding can be used for joining acrylic parts.

However, materials with different melting points are not good candidates for ultrasonic welding, since even a few degrees difference can result in one material melting before the other reaches its melting point, preventing a fusion between the parts.

Mechanical fastening, which concentrates loads at fastening points, is not recommended for acrylic parts, as the act of drilling holes or torquing fasteners can introduce potentially damaging stress. Holes should be cored out rather than drilled.

Acrylic Case History

Selecting the wrong grade of acrylic can lead to unsatisfactory results. Recently, a manufacturer of a blood-clot analyzer experienced early production problems with several complex parts molded from a general-purpose grade of acrylic.

With the formulation initially chosen, parts developed stress cracks upon ejection from the tools, primarily because of complicated geometries that included sharp corners, edges, and points.

Streaking also occurred. It was determined that additional toughness and impact strength were required, in a high-flow grade that would fill the molds.

An impact-modified, gamma- resistant acrylic--which is supplied in injection-grade pellets--was then selected and found to solve some tricky molding problems.

High tensile strength and impact resistance were critical; since the parts are used in the assembly of a cassette intended to contain blood, they must be resilient enough to resist cracking if accidentally dropped.

Dimensional precision was also important, given the complexity of the processes that occur under pressure inside the cassette during blood analysis.

A third requirement was UV transmissability, which was necessary for curing the adhesive used in the cassette assembly.

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Molded from the impact-modified grade, the cassette parts have proven to be sufficiently durable, both during fabrication and in use.

The acrylic material provides the requisite impact strength as well as a high degree of optical clarity and scratch resistance, and presents a pleasing, glossy surface finish.

After use in the blood-clot analyzer, the disposable cassettes are incinerated with other medical waste.

Disposal and Recycling

Acrylic burns extremely clean, providing virtually smokeless combustion with end products of carbon dioxide and water.

In addition, the material offers superior recyclability: acrylic can be reground and reused, which results in less material waste during molding.

Another characteristic of acrylic is its ability to be depolymerized back to its monomer, thoroughly purged of impurities, and repolymerized back into PMMA.

Commercial processing facilities set up for this process typically use a molten lead bath to vaporize the acrylic material.

The vapors are captured and recondensed into MMA, while eliminating any biomedical waste. This is a true recycling process, taking the polymer back to its monomer, whereas most other "recycling" processes involve crushing the material and using it in applications with lower specifications.

For environmental reasons, commercial molten lead baths are no longer operated in the United States, but do exist in England, India, and several other countries.

Used in medical applications for many years, acrylic offers a number of advantages compared with other clear thermoplastic materials traditionally specified for diagnostic equipment and medical devices.

Benefits include unsurpassed clarity, superior toughness, rapid gamma recovery, excellent UV-light transmittance, biocompatibility, and good chemical and scratch resistance.

Acrylic is easily processed and assembled and offers the potential to be recycled back to its monomer and used again, an important property as the pressure to recycle increases.

Acrylic resin can be custom formulated to meet the requirements of a broad range of medical applications. Since it can be processed in molds designed for polycarbonate and other thermoplastics, manufacturers using those materials should consider trial runs with acrylic.

Direct comparisons between parts molded with acrylic and with competing thermoplastics may reveal potential cost savings as well as unexpected side benefits such as improved surface finish, transparency, processing, and part performance.

Fluorocarbon polymers: Fluorocarbons are chemical compounds that contain carbon-fluorine bonds.

The relatively low reactivity and high polarity of the carbon-fluorine bond imparts unique characteristics to fluorocarbons.

Fluorocarbons tend to be only slowly broken down in the environment and therefore many are considered persistent organic pollutants. Many commercially useful fluorocarbons also contain hydrogen, chlorine, or bromine. Chlorofluorocarbons and hydrofluorocarbons

Chlorofluorocarbons (CFCs) are fluorocarbons that also contain chlorine atoms. They were formerly used widely in industry as refrigerants, propellants, and cleaning solvents (dichlorodifluoromethane and chlorodifluoromethane were among the most widely used refrigerants).

However, CFCs generally have potent ozone-depleting potential primarily due to homolytic cleavage of the carbon-chlorine bonds. Their use has now been mostly prohibited by the Montreal Protocol.

Hydrofluorocarbons (HFCs) are hydrocarbons in which some, but not all, of the hydrogen atoms have been replaced with fluorine.

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The fluorine atoms in these compounds do not catalyse ozone destruction, therefore HFCs do not damage the ozone layer. Consequently, HFCs such as tetrafluoroethane have become favored replacements for CFCs. Fluoropolymers

Fluorocarbon polymers are also well-known. These polymers are tough, chemical inert, and electrically insulating.

The most famous example is DuPont's Teflon, a polymer of the monomer tetra fluoro ethylene.

Other important polymers include polyvinylidene fluoride ([CH2CF2]n) and polychloro trifluoroethylene ([CFClCF2]n or PCTFE, or Kel-F). Uses Anesthetics

Many volatile anesthetics used to render surgical patients unconscious are fluorocarbons, such as methoxyflurane, enflurane, isoflurane, sevoflurane and desflurane.

The fluorine atoms reduce their flammability compared to the non-fluorinated anesthetics originally used, such as diethyl ether and cyclopropane, which are dangerously flammable. Solvents

Flurocarbons are used as industrial solvents due to their specific properties, including: non-flammability, stability, excellent dielectric properties, low surface tension and viscosity, very Main applications are:

• Precision Cleaning (Degreasing) • Electronic Assemblies Defluxing • Particulate Removal • Drying after Aqueous Cleaning • as a Carrier Fluid • as a Dielectric Coolant HFCs, particularly 1,1,1,2-tetrafluoroethane, are used for specialist extraction of extremely

important natural products; such as Taxol for cancer treatment from yew needles, evening primrose oil food supplement, and vanilla.

The use of 1,1,1,2-tetrafluroethane compliments other methods of extraction, in being highly selective and allowing high quality and high yield extractions.[1] Water repellant and stain repellant products

In general, highly fluorinated organic compounds are hydrophobic and have water-repellant and stain-repellant properties.

The original formulations of products such as Scotchgard contained fluorocarbons including perfluorobutane sulfonate and perfluorooctane sulfonate (PFOS).

But many of these uses have been phased out due to environmental concerns, such as those associated with perfluorooctanoic acid, an intermediate in the manufacture of PFOS. Similarly, products containing Gore-Tex and Teflon are made from fluoropolymers.

Fluorocarbons are also used in fishing line, in myriad precision plastics applications, and in highly precise lubrication applications. Biological role

Although there are thousands of known naturally-occurring organic compounds containing chlorine and bromine, there are only a handful of natural fluorocarbons.

They have been found in microorganisms and plants, but not animals. The most common natural fluorocarbon is fluoroacetic acid, a potent toxin found in a few species of plants.

Others included ω-fluoro fatty acids, fluoroacetone, and 2-fluorocitrate which are all believed to be biosynthesized from fluoroacetic acid.

Since the C-F bond is generally metabolically stable and fluorine is considered a bioisostere of the hydrogen atom, many pharmaceuticals contain C-F bonds.

An example of this is fluorinated uracil. When elemental fluorine is reacted with uracil, 5-fluorouracil is produced.

The resulting compound is an anticancer drug (antimetabolite) used to masquerade as uracil during the nucleic acid replication process.

This can lead to the incorporation of 5-fluorouracil into DNA and RNA as well as inhibition of the enzymes that are responsible for the synthesis of the normal components of DNA.

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These factors can be toxic to cancer cells that need to rapidly produce normal nucleic acids in order to continue growing.

Well known pharmaceutical drugs incorporating fluorine include fluoxetine (Prozac), paroxetine (Paxil), ciprofloxacin (Cipro), mefloquine, fluconazole, and many more. Chemical properties

The carbon-fluorine bond length is typically about 1.4 Å (1.39 Å in fluoromethane). This is shorter than any other carbon-halogen bond, and comparable in length to a carbon-hydrogen bond.

Since fluorine is a very electronegative atom (much more so than carbon), the carbon-fluorine bond has a significant dipole moment.

The carbon-fluorine bond is stronger than other carbon-halogen bonds. The bond dissociation energy is 552 kJ/mol for carbon-fluorine compared to 397, 288, 209 kJ/mol for bonds between carbon and chlorine, bromine and iodine, respectively.

The strength of the carbon-fluorine bond is also significantly stronger than the carbon-hydrogen bond, which is only 338 kJ/mol.

As a result of these unique features of the carbon-fluorine bond, an overarching theme in fluorocarbon chemistry is the contrasting set of physical and chemical properties in comparison to the corresponding hydrocarbons. Case studies follow. Pentakis(trifluoromethyl)cyclopentadiene

Pentakis(trifluoromethyl)cyclopentadiene (C5(CF3)5H) is a strong acid, with a pKa = −2. Its high acidity and robustness is indicated by the fact that this compound is typically purified by distillation from H2SO4.

In contrast, C5(CH3)5H requires a strong base such as butyllithium for deprotonation, as is typical for a hydrocarbon.[7] This compound is prepared in a multistep, one-pot reaction of potassium fluoride (KF) with 1,1,2,3,4,4-hexachlorobutadiene. Hexafluoroacetone and its imine

The molecule hexafluoroacetone ((CF3)2CO), the fluoro-analogue of acetone, has a boiling point of −27 °C compared to +55 °C for acetone itself.

This difference illustrates one of the remarkable effects of replacing C-H bonds with C-F bonds. Normally, the replacement of H atoms with heavier halogens results in elevated boiling points due to increased van der Waals interactions between molecules.

Further demonstrating the remarkable effects of fluorination, (CF3)2CO forms a stable, distillable hydrate, (CF3)2C(OH)2. Ketones rarely form stable hydrates.

Continuing this trend, (CF3)2CO adds ammonia to give (CF3)2C(OH)(NH2) which can be dehydrated with POCl3 to give (CF3)2CNH.[9] Compounds of the type R2C=NH are otherwise quite rare. Bulk polytetrafluoroethylene, also known as, e.g., PTFE, (CF.sub.2).sub.n, and Teflon.TM., is characterized by superior mechanical and electrical properties that are important for a wide range of applications.

For example, bulk PTFE is characterized by a low dielectric constant of about 2.1 and a low dielectric loss factor of less than about 0.0003 between about 60 Hz and 30,000 MHz.

Bulk PTFE is also characterized by high chemical stability, exemplified by its immunity to even strong alkalis and boiling hydrofluoric acid; low water absorption, exemplified by its water uptake of only about 0.005 weight % in a 24 hour period; and high thermal stability, exemplified by its weight loss of only about 0.05 weight % per hour at about 400.degree. C. A low coefficient of friction of between about 0.05-0.08 and a low permeability constant also characterize bulk PTFE. Bulk PTFE is also well-accepted as a substantially bio-compatible material that is tolerable by biological systems such as the human body.

Many biomedical and other applications are not optimally addressed by bulk PTFE, however.

For example, biologically-implantable devices such as neural probes, catheter inserts, implantable tubing, and other such devices, all of which are becoming increasingly complicated in geometry, are preferably encapsulated with a film to render the devices impervious to a biological environment, rather than being housed in a bulky PTFE package structure.

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Such implantable devices typically require of an encapsulating film not only the desired biological compatibility, but due to complex topology and connections to lead wires and associated circuitry, also inherently require an encapsulating film to be conformal and thin, as well as electrically insulating, tough, and flexible. Such a film should further be a good permeation barrier against the implantation environment.

A bulk PTFE package structure is thus not optimally applicable to such configurations.

There have been proposed various film deposition processes devised with the aim of producing thin films having properties similar to that of bulk PTFE.

For example, continuous radio-frequency plasma-enhanced chemical vapor deposition techniques have been proposed for producing PTFE-like films.

The films typically produced by such processes have been found, however, to be substantially lacking in one or more critical properties. In particular, the stoichiometry of the resulting films generally differs rather widely from that of bulk PTFE.

A typical ratio of fluorine to carbon (F/C ratio) for these films is only about 1.6, whereas bulk PTFE is characterized by a F/C ratio of 2.0. The films produced by various proposed processes are also typically characterized by a low fraction of CF.sub.2 groups; in contrast, bulk PTFE is composed substantially of CF.sub.2 groups. The high degree of crosslinking corresponding to low CF.sub.2 fractionality results in film brittleness, which is unacceptable for applications in which it is desired to encapsulate a flexible, bendable structure in a fluorocarbon film.

Furthermore, carbon-1s X-ray photoemission spectroscopy (XPS) of films deposited by various of the proposed processes reveals that in addition to requisite CF.sub.2 groups, comparable concentrations of unwanted moieties such as CF.sub.3, CF, and quaternary carbon moieties are also found in the deposited films.

Also, unlike bulk PTFE, films produced by the proposed processes typically contain carbon--carbon double bonds and further contain a significant concentration of dangling bonds. The unpaired electrons of these dangling bonds can be present in concentrations as high as about 10.sup.18 -10.sup.20 spins/cm.sup.3, and result in highly reactive film surface sites.

Taken together, these various unwanted moieties and defects present in typical encapsulating films result in film properties that are sharply degraded below that of bulk PTFE. For example, the dangling bonds in the films, being reactive sites, can react with atmospheric oxygen or water, producing aging effects that result in undesirable film property variations over time.

The dielectric properties of the films are also degraded, resulting in, e.g., an increased dielectric loss. The chemical and thermal stability of the films are also degraded, due to the suboptimal film stoichiometry.

Thus, deposition techniques devised in an effort to duplicate properties of bulk PTFE for thin-film applications have generally resulted in only suboptimal thin films that typically cannot adequately address performance requirements of PTFE thin film applications. Rubbers:

Natural rubber contains a group of water soluble or extractable proteins (sometimes referred to as “EP”) that can cause allergic reactions in the human body when contact is made over a prolonged period of time.

The terminology of “Latex-Free" is often used to mean not made from natural rubber (latex) or not containing natural rubber. Often the term “synthetic bands” is used to refer to Latex-Free bands. The bottom line is that the absence of latex eliminates the risk of latex protein allergy.

Silicone, natural, and synthetic rubbers have been used for the fabrication of implants. Natural rubber is made mostly from the latex of the Hevea brasiliensis tree and the chemical formula is the same as that of cis-1,4 polyisoprene.

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Natural rubber was found to be compatible with blood in its pure form. Also, cross-linking by X-ray and organic peroxides produces rubber with superior blood compatibility compared with rubbers made by the conventional sulfur vulcanization.

Synthetic rubbers were developed to substitute for natural rubber. The Ziegler-Natta types of stereospecific polymerization techniques have made this variety possible.

The synthetic rubbers have rarely been used to make implants. The physical properties vary widely due to the wide variations in preparation recipes of these rubbers.

Silicone rubber, developed by Dow Corning company, is one of the few polymers developed for medical use. The repeating unit is dimethyl siloxane which is polymerized by a condensation polymerization.

Low molecular weight polymers have low viscosity and can be cross-linked to make a higher molecular weight, rubber-like material. Medical grade silicone rubbers contain stannous octate as a catalyst and can be mixed with a base polymer at the time of implant fabrication. Thermoplastics: Polyurethanes:

Polyurethanes are usually thermosetting polymers: they are widely used to coat implants. Polyurethane rubbers are produced by reacting a prepared prepolymer chain with an aromatic di-isocyanate to make very long chains possessing active isocyanate groups for cross-linking.

The polyurethane rubber is quite strong and has good resistance to oil and chemicals. Polyacetal, Polysulfone, and Polycarbonate:

These polymers have excellent mechanical, thermal, and chemical properties due to their stiffened main backbone chains.

Polyacetals and polysulfones are being tested as implant materials, while polycarbonates have found their applications in the heart/lung assist devices, food packaging, etc.

Polyacetals are produced by reacting formaldehyde. These are also sometimes called polyoxymethylene (POM) and known widely as Delrin®

(DuPont). These polymers have a reasonably high molecular weight (>2 × 104 g/mol) and have

excellent mechanical properties. More importantly, they display an excellent resistance to most chemicals and to water over

wide temperature ranges. Polysulfones were developed by Union Carbide in the 1960s. These polymers have a high

thermal stability due to the bulky side groups (therefore, they are amorphous) and rigid main backbone chains.

They are also highly stable to most chemicals but are not so stable in the presence of polar organic solvents such as ketones and chlorinated hydrocarbons.

Polycarbonates are tough, amorphous, and transparent polymers made by reacting bisphenol A and diphenyl carbonate.

It is noted for its excellent mechanical and thermal properties (high Tg:150°C), hydrophobicity, and antioxidative properties. Physiochemical characteristics of biopolymers:

A charged biomaterial including at least one first hydrophilic polymer, and super paramagnetic iron oxide particles complexed with a second hydrophilic polymer substantially identical to or different from the first hydrophilic polymer, wherein the superparamagnetic iron oxide particles complexed with the second hydrophilic polymer are distributed substantially homogeneously and substantially without aggregates in the first hydrophilic polymer.

A process for preparing a charged biomaterial including forming an aqueous solution of hydrophilic monomers and superparamagnetic iron oxide particles complexed with a hydrophilic polymer including monomers substantially identical to or different from the hydrophilic monomers, polymerizing the solution and forming a polymer hydrogel in which are distributed substantially homogeneously and substantially without aggregates the superparamagnetic iron oxide particles complexed with the hydrophilic polymer. Biodegradable polymers for medical purposes:

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Many polymers have been used in biomedical applications, including polyesters, polyvinylacetate, polyacrylates, polyorthoesters, poly(amino acids) and polyanhydrides. Biodegradable polymers, such as polyesters and polyanhydrides, particularly those that are biodegradable into natural metabolites, have been the subject of significant research efforts directed toward the development of new multifunctional polymer compositions for implant fabrication.

The novel multifunctional biodegradable polymers are prepared for use in construction/formulation of implants useful for tissue treatment and repair.

The biodegradable polymers are amphiphilic in that they include a natural hydrophobic or lipophilic group conjugated through a divalent linker, preferably a polymer biodegradable into natural metabolites, to a hydrophilic moiety which exhibits affinity for natural tissues.

Preferred embodiments of the novel biodegradable polymers in accordance with this invention exhibit the property of self-assembly, detectable by x-ray detraction techniques, characteristic of art-recognized liquid crystal compositions.

The present biodegradable polymers can be used in formulating implant matrix compositions for tissue treatment and/or repair wherein the polymer provides a biodegradable carrier matrix for added biofunctional adjuvants.

The implant matrix compositions can be in the form of a flowable viscous liquid or gel, a paste, a moldable putty, or a shape-retaining solid in molded form, depending on the molecular weight and chemical constituents of the component amphiphilic polymer(s).

The cooperation of physical and chemical properties of the present amphiphilic biodegradable polymers work to produce a unique implant matrix functionality.

The self-assembling character of preferred polymers in accordance with this invention can impart a unique matrix structure with localized ordered domains.

That property, coupled with the unique molecular structure of the present biodegradable polymers, provides significant potential for fabrication of implants with improved functionality.

Moreover, the liquid crystalline character of preferred biodegradable polymers in accordance with this invention enables the fabrication of implant matrix compositions that exhibit pre-determined temperature dependent order allowing, for example, the preparation of an implantable matrix composition that exhibits fluid (amorphous) properties at a temperature slightly above body temperature and solid properties at body temperature.

An implant matrix composition so formulated can be injected into the locus of a tissue defect as a liquid and thereafter solidify with phase transition to a microcrystalline or liquid crystalline form at body temperature.

The present novel biodegradable polymers, when used in implant matrix compositions degrade (hydrolyze) with time after implantation, thereby providing a temporary template for cell (of endogenous or exogenous origin) growth and release of other contained bioactive adjuvants, if any, to the locus of the implant.

When used for tissue repair the implant matrix is gradually replaced by endogenous tissue in repair of a targeted tissue defect. Biodegradable Materials in Tissue Engineering and Regenerative Medicine

The term biodegradation is loosely associated with materials that could be broken down by nature either through hydrolytic mechanisms without the help of enzymes and/or enzymatic mechanism. Other terms like absorbable, erodible, and resorbable have also been used in the literature to indicate biodegradation.

This is because this class of biomaterials has two major advantages that non-biodegradable biomaterials do not have.

First, they do not elicit permanent chronic foreignbody reactions due to the fact that they are gradually absorbed by the human body and do not permanently leave traces of residual in the implantation sites.

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Second, some of them have recently been found to be able to regenerate tissues, so called tissue engineering , through the interaction of their biodegradation with immunologic cells like macrophages. Hence, surgical implants made from biodegradable biomaterials could be used as a temporary scaffold for tissue regeneration. Properties of Commercially Important Synthetic Absorable Polymers: .

Fiber . Strength (MPa) Modulus (GPa ) Elongation (%)

Polymer Crystallinity . PGA High 890 8.4 30 PLLA High 900 8.5 25 PLA None — — — Polyglactin910 High 850 8.6 24 Polydioxanone High 490 2.1 35 Polyglyconate High 550 2.4 45 Poliglecaprone25 — 91,100 113,000 39 .

The earliest and most commercially significant biodegradable polymeric biomaterials were originated from linear aliphatic polyesters like polyglycolide and polylactide from poly (α-hydroxyacetic acids)

The recent introduction of several new synthetic and natural biodegradable polymeric biomaterials extends the domain beyond this family of simple polyesters.

These new ommercially significant biodegradable polymeric biomaterials include i poly(orthoesters) ii polyanhydrides, iii polysaccharides iv. poly(esteramides) v. Tyrosine-based polyarylates or polyiminocarbonates or polycarbonates, vi. Poly(D,L-lactide-urethane) vii. Poly(β -hydroxybutyrate) viii. Poly(ε -caprolactone), ix. Poly[bis(carboxylatophenoxy) x. Phosphazene] xi. Poly(amino acids) xii. Pseudo-poly(amino acids) xiii. Copolymers derived from amino acids and non-amino acids.

All the above biodegradable polymeric biomaterials could be generally divided into eight groups based on their chemical origin:

(1) Biodegradable linear aliphatic polyesters (e.g., polyglycolide, polylactide, polycaprolactone, polyhydroxybutyrate) and their copolymers within the aliphatic polyester family like poly(glycolide-L-lactide) copolymer and poly(glycolide- ε -caprolactone) copolymer; (2) Biodegradable copolymers between linear aliphatic polyesters in (1) and monomers other than linear aliphatic polyesters like, poly(glycolide-trimethylene carbonate) copolymer, poly(L-lactic acid-L-lysine) copolymer, Tyrosinebased polyarylates or polyiminocarbonates or polycarbonates, poly(D,L-lactide-urethane), and poly(esteramide) (3) Polyanhydrides (4) Poly(orthoesters) (5) Poly(ester-ethers) like poly-p-dioxanone (6) Biodegradable polysaccharides like hyaluronic acid, chitin and chitson (7) polyamino acids like poly-Lglutamic acid and poly-L-lysine (8) Inorganic biodegradable polymers like polyphosphazene and poly

[bis(carboxylatophenoxy)phosphazene]

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The most successful, and frequent biomedical applications of biodegradable polymeric biomaterials have been in wound closure

All biodegradable wound closure biomaterials are based upon the glycolide and lactide families. For example, polyglycolide, poly(glycolide-L-lactide), poly(ester-ether), poly(glycolide-trimethylene carbonate) random block copolymer, and poly(glycolide- ε -caprolactone) copolymer.

The next largest biomedical application of biodegradable polymeric biomaterials that are commercially satisfactory is drug control/release devices.

Besides the commercial uses the biodegradable polymeric biomaterials have been experimented with as (1) vascular grafts, (2) vascular stents, (3) vascular couplers for vessel anastomosis, (4) nerve growth conduits, (5) augmentation of defected bone, (6) ligament/tendon prostheses,(7) intramedullary plug during total hip replacement, (8) anastomosis ring for intestinal surgery, and (9) stents in ureteroureterostomies for accurate suture placement.

Synthetic Biodegradable Polymers as Medical Devices

In the first half of this century, research into materials synthesized from glycolic acid and other -hydroxy acids was abandoned for further development because the resulting polymers were too unstable for long-term industrial uses.

However, this very instability—leading to biodegradation—has proven to be immensely important in medical applications over the last three decades.

Polymers prepared from glycolic acid and lactic acid have found a multitude of uses in the medical industry, beginning with the biodegradable sutures first approved in the 1960s.

Since that time, diverse products based on lactic and glycolic acid—and on other materials, including poly(dioxanone), poly(trimethylene carbonate) copolymers, and poly (-caprolactone) homopolymers and copolymers—have been accepted for use as medical devices.

In addition to these approved devices, a great deal of research continues on polyanhydrides, polyorthoesters, polyphosphazenes, and other biodegradable polymers.

A biodegradable intravascular stent prototype is molded from a blend of polylactide and trimethylene carbonate. Why would a medical practitioner want a material to degrade?

There may be a variety of reasons, but the most basic begins with the physician's simple desire to have a device that can be used as an implant and will not require a second surgical intervention for removal.

Besides eliminating the need for a second surgery, the biodegradation may offer other advantages.

For example, a fractured bone that has been fixated with a rigid, nonbiodegradable stainless implant has a tendency for refracture upon removal of the implant.

Because the stress is borne by the rigid stainless steel, the bone has not been able to carry sufficient load during the healing process.

However, an implant prepared from biodegradable polymer can be engineered to degrade at a rate that will slowly transfer load to the healing bone.

Another exciting use for which biodegradable polymers offer tremendous potential is as the basis for drug delivery, either as a drug delivery system alone or in conjunction to functioning as a medical device.

Polymer scientists, working closely with those in the device and medical fields, have made tremendous advances over the last 30 years. This article will focus on a number of these developments. We will also review the chemistry of the polymers, including synthesis and degradation, describe how properties can be controlled by proper synthetic controls such as copolymer composition, highlight special requirements for processing and handling, and discuss some of the commercial devices based on these materials. POLYMER CHEMISTRY

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Biodegradable polymers can be either natural or synthetic. In general, synthetic polymers offer greater advantages than natural materials in that they can be tailored to give a wider range of properties and more predictable lot-to-lot uniformity than can materials from natural sources. Synthetic polymers also represent a more reliable source of raw materials, one free from concerns of immunogenicity.

Table I. Properties of common biodegradable polymers.

The general criteria for selecting a polymer for use as a biomaterial is to match the mechanical properties and the time of degradation to the needs of the application (see Table I). The ideal polymer for a particular application would be configured so that it:

• Has mechanical properties that match the application, remaining sufficiently strong until the surrounding tissue has healed.

• Does not invoke an inflammatory or toxic response. • Is metabolized in the body after fulfilling its purpose, leaving no trace. • Is easily processable into the final product form. • Demonstrates acceptable shelf life. • Is easily sterilized. The factors affecting the mechanical performance of biodegradable polymers are those that

are well known to the polymer scientist, and include monomer selection, initiator selection, process conditions, and the presence of additives.

These factors in turn influence the polymer's hydrophilicity, crystallinity, melt and glass-transition temperatures, molecular weight, molecular-weight distribution, end groups, sequence distribution (random versus blocky), and presence of residual monomer or additives.

In addition, the polymer scientist working with biodegradable materials must evaluate each of these variables for its effect on biodegradation.

Biodegradation has been accomplished by synthesizing polymers that have hydrolytically unstable linkages in the backbone. The most common chemical functional groups with this characteristic are esters, anhydrides, orthoesters, and amides. We will discuss the importance of the properties affecting biodegradation later in the article.

Polymer Melting Point (°C)

Glass-Transition Temp (°C)

Modulus (Gpa)a

Degradation Time (months)b

PGA 225—230 35—40 7.0 6 to 12 LPLA 173—178 60—65 2.7 >24 DLPLA Amorphous 55—60 1.9 12 to 16 PCL 58—63 (—65)— (—60) 0.4 >24

PDO N/A (—10)— 0 1.5 6 to 12 PGA-TMC N/A N/A 2.4 6 to 12

85/15 DLPLG

Amorphous 50—55 2.0 5 to 6

75/25 DLPLG

Amorphous 50—55 2.0 4 to 5

65/35 DLPLG

Amorphous 45—50 2.0 3 to 4

50/50 DLPLG

Amorphous 45—50 2.0 1 to 2

a Tensile or flexural modulus. b Time to complete mass loss. Rate also depends on part geometry.

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The synthetic biodegradable polymers that are currently being used or investigated for use in wound closure:

The synthetic biodegradable polymers that are currently being used or investigated for use in wound closure (sutures, staples); orthopedic fixation devices (pins, rods, screws, tacks, ligaments); dental applications (guided tissue regeneration); cardiovascular applications (stents, grafts); and intestinal applications (anastomosis rings).

Most of the commercially available biodegradable devices are polyesters composed of homopolymers or copolymers of glycolide and lactide.

There are also devices made from copolymers of trimethylene carbonate and -caprolactone, and a suture product made from polydioxanone. Polyglycolide (PGA).

Polyglycolide is the simplest linear aliphatic polyester. PGA was used to develop the first totally synthetic absorbable suture.

Glycolide monomer is synthesized from the dimerization of glycolic acid. Ring-opening polymerization yields high-molecular-weight materials, with approximately 1—3% residual monomer present (see Figure 1).

PGA is highly crystalline (45—55%), with a high melting point (220—225°C) and a glass-transition temperature of 35—40°C. Because of its high degree of crystallization, it is not soluble in most organic solvents; the exceptions are highly fluorinated organics such as exafluoroisopropanol.

Fibers from PGA exhibit high strength and modulus and are too stiff to be used as sutures except in the form of braided material. Sutures of PGA lose about 50% of their strength after 2 weeks and 100% at 4 weeks, and are completely absorbed in 4—6 months.

Glycolide has been copolymerized with other monomers to reduce the stiffness of the resulting fibers. Polylactide (PLA).

Lactide is the cyclic dimer of lactic acid that exists as two optical isomers, d and l. l-lactide is the naturally occurring isomer, and dl-lactide is the synthetic blend of d-lactide and l-lactide. The homopolymer of l-lactide (LPLA) is a semicrystalline polymer.

These types of materials exhibit high tensile strength and low elongation, and consequently have a high modulus that makes them more suitable for load-bearing applications such as in orthopedic fixation and sutures.

Poly(dl-lactide) (DLPLA) is an amorphous polymer exhibiting a random distribution of both isomeric forms of lactic acid, and accordingly is unable to arrange into an organized crystalline structure.

This material has lower tensile strength, higher elongation, and a much more rapid degradation time, making it more attractive as a drug delivery system.

Poly(l-lactide) is about 37% crystalline, with a melting point of 175—178°C and a glass-transition temperature of 60—65°C.

The degradation time of LPLA is much slower than that of DLPLA, requiring more than 2 years to be completely absorbed. Copolymers of l-lactide and dl-lactide have been prepared to disrupt the crystallinity of l-lactide and accelerate the degradation process.

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Poly( -caprolactone).

The ring-opening polymerization of -caprolactone yields a semicrystalline polymer with a melting point of 59—64°C and a glass-transition temperature of —60°C (see Figure 2).

The polymer has been regarded as tissue compatible and used as a biodegradable suture in Europe. Because the homopolymer has a degradation time on the order of 2 years, copolymers have been synthesized to accelerate the rate of bioabsorption.

For example, copolymers of -caprolactone with dl-lactide have yielded materials with more-rapid degradation rates. A block copolymer of -caprolactone with glycolide, offering reduced stiffness compared with pure PGA, is being sold as a monofilament suture

Poly( -caprolactone). Poly(dioxanone) (a polyether-ester).

The ring-opening polymerization of p-dioxanone (see Figure 3) resulted in the first clinically tested monofilament synthetic suture, known as PDS (marketed by Ethicon).

This material has approximately 55% crystallinity, with a glass-transition temperature of —10 to 0°C. The polymer should be processed at the lowest possible temperature to prevent depolymerization back to monomer.

Poly(dioxanone) has demonstrated no acute or toxic effects on implantation. The monofilament loses 50% of its initial breaking strength after 3 weeks and is absorbed

within 6 months, providing an advantage over Dexon or other products for slow-healing wounds. Poly(lactide-co-glycolide).

Using the polyglycolide and poly(l-lactide) properties as a starting point, it is possible to copolymerize the two monomers to extend the range of homopolymer properties.

Copolymers of glycolide with both l-lactide and dl-lactide have been developed for both device and drug delivery applications.

It is important to note that there is not a linear relationship between the copolymer composition and the mechanical and degradation properties of the materials.

For example, a copolymer of 50% glycolide and 50% dl-lactide degrades faster than either homopolymer .

Copolymers of l-lactide with 25—70% glycolide are amorphous due to the disruption of the regularity of the polymer chain by the other monomer.

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A copolymer of 90% glycolide and 10% l-lactide was developed by Ethicon as an absorbable suture material under the trade name Vicryl. It absorbs within 3—4 months but has a slightly longer strength-retention time.

Copolymers of glycolide with trimethylene carbonate (TMC), called have been prepared as

both sutures and as tacks and screws. Typically, these are prepared as A-B-A block copolymers in a 2:1 glycolide:TMC ratio,

with a glycolide-TMC center block (B) and pure glycolide end blocks (A). These materials have better flexibility than pure PGA and are absorbed in approximately 7

months. Glycolide has also been polymerized with TMC and p-dioxanone (Biosyn, by United States Surgical Corp., Norwalk, CT) to form a terpolymer suture that absorbs within 3—4 months and offers reduced stiffness compared with pure PGA fibers. Other Polymers under Development.

Currently, only devices made from homopolymers or copolymers of glycolide, lactide, caprolactone, p-dioxanone, and trimethylene carbonate have been cleared for marketing by FDA. A number of other polymers, however, are being investigated for use as materials for biodegradable devices.

In addition to their suitability for medical uses, biodegradable polymers make excellent candidates for packaging and other consumer applications.

A number of companies are evaluating ways to make low-cost biodegradable polymers. One method is to bioengineer the synthesis of the polymers, using microorganisms to produce energy-storing polyesters.

Two examples of these materials—polyhydroxybutyrate (PHB) and polyhydroxyvalerate (PHV)—are commercially available as copolymers under the trade name Biopol (Monsanto Co., St. Louis) and have been studied for use in medical devices.

The PHB homopolymer is crystalline and brittle, whereas the copolymers of PHB with PHV are less crystalline, more flexible, and easier to process.

These polymers typically require the presence of enzymes for biodegradation but can degrade in a range of environments and are under consideration for several biomedical applications.

The use of synthetic poly(amino acids) as polymers for biomedical devices would seem a logical choice, given their wide occurrence in nature.

In practice, however, pure insoluble poly(amino acids) have found little utility because of their high crystallinity, which makes them difficult to process and results in relatively slow degradation.

The antigenicity of polymers with more than three amino acids in the chain also makes them inappropriate for use in vivo.

To circumvent these problems, modified "pseudo" poly(amino acids) have been synthesized by using a tyrosine derivative.

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Tyrosine-derived polycarbonates, for example, are high-strength materials that may be useful as orthopedic implants.

It is also possible to copolymerize poly(amino acids) to modify their properties. The group that has been researched most extensively is the polyesteramides. PACKAGING AND STERILIZATION

Because biodegradable polymers are hydrolytically unstable, the presence of moisture can degrade them in storage, during processing, and after device fabrication

The as-synthesized polymers have relatively low water contents, since any residual water in the monomer is used up in the polymerization reaction. The polymers are quickly packaged after manufacture—generally double-bagged under an inert atmosphere or vacuum.

The bag material may be polymeric or foil, but it must be highly resistant to water permeability.

To minimize the effects of any moisture present, the polymers are typically stored in a freezer. Packaged polymers should always be at room temperature when opened to minimize condensation, and should be handled as little as possible at ambient atmospheric conditions.

As expected, there is a relationship among biodegradation rate, shelf stability, and polymer properties. For instance, the more hydrophilic glycolide polymers are much more sensitive to hydrolytic degradation than are polymers prepared from the more hydrophobic lactide.

Final packaging consists of placing the suture or device in an airtight, moistureproof container.

A desiccant can be added to further reduce the effects of moisture. Sutures, for example, are wrapped around a specially dried paper holder that acts as a desiccant. In some cases, the finished device may be stored at subambient temperature as an added precaution against degradation. Devices incorporating biodegradable polymers cannot be subjected to autoclaving, and must be sterilized by gamma or E-beam irradiation or by exposure to ethylene oxide (EtO) gas.

There are certain disadvantages, however, to both irradiation and EtO sterilization. Irradiation, particularly at doses above 2 Mrd, can induce significant degradation of the polymer chain, resulting in reduced molecular weight as well as influencing final mechanical properties and degradation times.

Polyglycolide, poly(lactide), and poly(dioxanone) are especially sensitive to ionizing radiation, and these materials are usually sterilized by EtO for device applications.

Because the highly toxic EtO can present a safety hazard, great care must be taken to ensure that all the gas is removed from the device before final packaging.

The temperature and humidity conditions should also be considered when submitting devices for sterilization.

Temperatures must be kept below the glass-transition temperature of the polymer to prevent the part geometry from changing during sterilization. If necessary, parts can be kept at 0°C or lower during the irradiation process. PROCESSING

All commercially available biodegradable polymers can be melt processed by conventional means such as injection molding, compression molding, and extrusion.

As with packaging, special consideration needs to be given to the exclusion of moisture from the material before melt processing to prevent hydrolytic degradation.

Special care must be taken to dry the polymers before processing and to rigorously exclude humidity during processing.

Because most biodegradable polymers have been synthesized by ring-opening polymerization, a thermodynamic equilibrium exists between the forward or polymerization reaction and the reverse reaction that results in monomer formation.

Excessively high processing temperatures may result in monomer formation during the molding or extrusion process.

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The presence of excess monomer can act as a plasticizer, changing the material's mechanical properties, and can catalyze the hydrolysis of the device, thus altering degradation kinetics. Therefore, these materials should be processed at the lowest temperatures possible.

DEGRADATION

Once implanted, a biodegradable device should maintain its mechanical properties until it is no longer needed and then be absorbed and excreted by the body, leaving no trace.

Simple chemical hydrolysis of the hydrolytically unstable backbone is the prevailing mechanism for the polymer's degradation.

This occurs in two phases. In the first phase, water penetrates the bulk of the device, preferentially attacking the chemical bonds in the amorphous phase and converting long polymer chains into shorter water-soluble fragments.

Because this occurs in the amorphous phase initially, there is a reduction in molecular weight without a loss in physical properties, since the device matrix is still held together by the crystalline regions.

The reduction in molecular weight is soon followed by a reduction in physical properties, as water begins to fragment the device.

In the second phase, enzymatic attack and metabolization of the fragments occurs, resulting in a rapid loss of polymer mass. This type of degradation—when the rate at which water penetrates the device exceeds that at which the polymer is converted into water-soluble materials (resulting in erosion throughout the device)—is called bulk erosion. All of the commercially available synthetic devices and sutures degrade by bulk erosion.

A second type of biodegradation, known as surface erosion, occurs when the rate at which the polymer penetrates the device is slower than the rate of conversion of the polymer into water-soluble materials.

Surface erosion results in the device thinning over time while maintaining its bulk integrity. Polyanhydrides and polyorthoesters are examples of materials that undergo this type of erosion—when the polymer is hydrophobic, but the chemical bonds are highly susceptible to hydrolysis.

In general, this process is referred to in the literature as bioerosion rather than biodegradation.

• The degradation-absorption mechanism is the result of many interrelated factors, including: • The chemical stability of the polymer backbone. • The presence of catalysts, additives, impurities, or plasticizers. • The geometry of the device. Balancing these factors by tailoring an implant to slowly degrade and transfer stress at the appropriate rate to surrounding tissues as they heal is one of the major challenges facing researchers today.

Factors That Accelerate Polymer Degradation • More hydrophilic backbone. • More hydrophilic endgroups. • More reactive hydrolytic groups in the backbone. • Less crystallinity. • More porosity. • Smaller device size.

COMMERCIAL BIODEGRADABLE DEVICES Sutures.

They are divided into braided and monofilament categories. Braided sutures are typically more pliable than monofilament and exhibit better knot security when the same type of knot is used. Monofilament sutures are more wiry and may require a more secure knot.

Their major advantage is that they exhibit less tissue drag, a characteristic that is especially important for cardiovascular, ophthalmic, and neurological surgeries.

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A recent source in the literature lists eight objective and three subjective parameters for suture selection based on criteria such as tensile strength, strength retention, knot security, tissue drag, infection potential, and ease of tying. Dental Devices.

Biodegradable polymers have found use in two dental applications. Employed as a void filler following tooth extraction, porous polymer particles can be packed into the cavity to aid in quicker healing. As a guided-tissue-regeneration (GTR) membrane, films of biodegradable polymer can be positioned to exclude epithelial migration following periodontal surgery.

The exclusion of epithelial cells allows the supporting, slower-growing tissue—including connective and ligament cells—to proliferate.

Three examples of these GTR materials are Resolut from W.L. Gore (Flagstaff, AZ), Atrisorb from Atrix Laboratories (Fort Collins, CO), and Vicryl mesh from Ethicon. Orthopedic Fixation Devices.

Orthopedic fixation devices made from synthetic biodegradable polymers have advantages over metal implants in that they transfer stress over time to the damaged area, allowing healing of the tissues, and eliminate the need for a subsequent operation for implant removal.

The currently available materials have not exhibited sufficient stiffness to be used as bone plates for support of long bones, such as the femur.

Rather, they have found applications where lower-strength materials are sufficient: for example, as interference screws in the ankle, knee, and hand areas; as tacks and pins for ligament attachment and meniscal repair; as suture anchors; and as rods and pins for fracture fixation.

Screws and plates of poly(l-lactide-co-glycolide) for craniomaxillo facial repair Other Applications.

Biodegradable polymers have found other applications that have been commercialized or are under investigation.

Anastomosis rings have been developed as an alternative to suturing for intestinal resection. Tissue staples have also replaced sutures in certain procedures. Other applications currently under scrutiny include ligating clips, vascular grafts, stents, and tissue-engineering scaffolds. Biodegradable Polymers in Tissue Engineering

One of the exciting current areas for applications of biodegradable polymers is in tissue engineering. Several companies are investigating using these materials as a matrix for living cells.

Important properties in this regard include porosity for cell in-growth, a surface that balances hydrophilicity and hydrophobicity for cellular attachment, mechanical properties that are compatible with those of the tissue, and degradation rate and by-products.

The polymer matrix may represent the device itself, or can be a scaffold for cell growth in vitro that is degraded by the growing cells prior to implantation.

The device can also be formulated to contain additives or active agents for more rapid tissue growth or compatibility. For example, a bone implant may contain a form of calcium phosphate or a growth factor such as one of the bone morphogenetic proteins.

There are a number of ways of making the three-dimensional matrices required for tissue engineering.

These methods include woven or nonwoven preparations from spun fibers, blown films using solvents or propellants, or sintered polymer particles.

One of the newest methods is being developed by Therics (Princeton, NJ), which has licensed a system for building three-dimensional devices for use as scaffolds and in drug delivery products. In this system, small spheres of polymer are laid out in thin films.

Using technology similar to that found in ink-jet printers, small amounts of solvent are used to fuse particles together.

The particles not fused are removed and another layer of particles laid out.

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This particle placement and fusing is continued for many layers, until the exact three-dimensional structure is obtained. Because each polymer layer is applied in a separate step, different polymers can be used to obtain different properties in the interior and exterior of the device. Synthetic polymeric membranes and their biological applications:

Novel multifunctional biodegradable polymers are prepared for use in onstruction /formulation of implants useful for tissue treatment and repair.

The biodegradable polymers are amphiphilic in that they include a natural hydrophobic or lipophilic group conjugated through a divalent linker, preferably a polymer biodegradable into natural metabolites, to a hydrophilic moiety which exhibits affinity for natural tissues.

Preferred embodiments of the novel biodegradable polymers in accordance with this invention exhibit the property of self-assembly, detectable by x-ray detraction techniques, characteristic of art-recognized liquid crystal compositions.

The present biodegradable polymers can be used in formulating implant matrix compositions for tissue treatment and/or repair wherein the polymer provides a biodegradable carrier matrix for added biofunctional adjuvants.

The implant matrix compositions can be in the form of a flowable viscous liquid or gel, a paste, a moldable putty, or a shape-retaining solid in molded form, depending on the molecular weight and chemical constituents of the component amphiphilic polymer(s).

The cooperation of physical and chemical properties of the present amphiphilic biodegradable polymers work to produce a unique implant matrix functionality. The self-assembling character of preferred polymers in accordance with this invention can impart a unique matrix structure with localized ordered domains.

That property, coupled with the unique molecular structure of the present biodegradable polymers, provides significant potential for fabrication of implants with improved functionality.

Moreover, the liquid crystalline character of preferred biodegradable polymers in accordance with this invention enables the fabrication of implant matrix compositions that exhibit pre-determined temperature dependent order allowing, for example, the preparation of an implantable matrix composition that exhibits fluid (amorphous) properties at a temperature slightly above body temperature and solid properties at body temperature. An implant matrix composition so formulated can be injected into the locus of a tissue defect as a liquid and thereafter solidify with phase transition to a microcrystalline or liquid crystalline form at body temperature.

The present novel biodegradable polymers, when used in implant matrix compositions degrade (hydrolyze) with time after implantation, thereby providing a temporary template for cell (of endogenous or exogenous origin) growth and release of other contained bioactive adjuvants, if any, to the locus of the implant.

When used for tissue repair the implant matrix is gradually replaced by endogenous tissue in repair of a targeted tissue defect. Biopolymers in controlled release systems:

Controlled release of biologically active compounds in the context of drug and vaccine delivery is an important area of research with broad implications in many areas of medicine. Controlled release of biologically active compounds in the context of drug and vaccine delivery is an important area of research with broad implications in many areas of medicine.

The emulsan family of biopolymers also has significant potential in vaccine and drug delivery based on recent studies. Each of these biopolymers offers exciting opportunities to modulate biological responses via control of chemistry and physical properties achieved during biosynthesis or postsynthesis modifications.

1. Polymers as Drug Delivery Carriers • Hydrophilic Cellulose Derivatives as Drug Delivery Carriers: Influence of Substitution

Type on the Properties of Compressed Matrix Tablets • Poly(Vinyl Alcohol) as a Drug Delivery Carrier

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• Development of Acrylate and Methacrylate Polymer Networks for Controlled Release by Photopolymerization Technology

• Smart Polymers for Controlled Drug Delivery • Complexing Polymers in Drug Delivery • Polylactic and Polyglycolic Acids as Drug Delivery Carriers • Use of Infrared and Raman Spectroscopy for Characterization of Controlled Release

Systems • Accurate Models in Controlled Drug Delivery Systems

2. Mechanism-Based Classification of Controlled Release Devices • Drug Release from Swelling-Controlled Systems • Superporous Hydrogels as a Platform for Oral Controlled Drug Delivery • Osmotic Implantable Delivery Systems • Bioadhesive Controlled Release Systems

3. Micro- and Nanoparticulate Release Systems • Microencapsulation Technology: Interfacial Polymerization Method • Nanoparticulate Controlled Release Systems for Cancer Therapy • Microencapsulation Using Coacervation/Phase Separation: An Overview of the

Technique and Applications • Microsphere Preparation by Solvent Evaporation Method • Nanosuspensions—A Formulation Approach for Poorly Soluble and Poorly

Bioavailable Drugs • Large-Scale Production of Solid Lipid Nanoparticles (SLN) and Nanosuspensions

(DissoCubes) • Solid Lipid Nanoparticles (SLN) as a Carrier System for the Controlled Release of

Drugs • Stability of Encapsulated Substances in Poly(Lactide-co-Glycolide) Delivery Systems • Development of Polysaccharides as Novel Drug Carrier Systems 4. Classification of Controlled Release Devices According to Administration Site • An Overview of Controlled Release Systems • Research and Development Aspects of Oral Controlled-Release Dosage Forms • A Gastrointestinal Retentive Microparticulate System to Improve Oral Drug Delivery • In Vitro-In Vivo Correlations in the Development of Solid Oral Controlled Release

Dosage Forms • Gamma Scintigraphy in the Analysis of the Behavior of Controlled Release Systems • Electrically Assisted Transdermal Delivery of Drugs • A Novel Method Based on Artificial Neural Networks for Optimizing Transdermal

Drug Delivery Systems • Transdermal Drug Delivery by Skin Electroporation • Enhancement of Transdermal Transport Using Ultrasound in Combination with Other

Enhancers • Electrotransport Systems for Transdermal Delivery: A Practical Implementation of

Iontophoresis 5. Peptide and Protein Release Systems • Controlled Release Protein Therapeutics: Effects of Process and Formulation on

Stability • Solid-State Chemical Stability of Peptides and Proteins: Application to Controlled

Release Formulations • Growth Factor Release from Biodegradable Hydrogels to Induce Neovascularization • Biopolymers for Release of Interleukin-2 for Treatment of Cancer 6. Medical Applications of Drug Delivery • Osmotic Drug Delivery from Asymmetric Membrane Film-Coated Dosage Forms • Controlled Release Pain Management Systems • Biodegradable Systems for Long-Acting Nestorone

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• Preparation and Evaluation of Buprenorphine Microspheres for Parenteral Administration

• Prolonged Release of Hydromorphone from a Novel Poly(Lactic-co-Glycolic) Acid Depot System: Initial In Vitro and In Vivo Observations

• Incorporation of an Active Agent into a Biodegradable Cement: Encapsulation of the Agent as Protection from Chemical Degradation During Cure and Effect on Release Profile

• The Pharmacoeconomic Value of Controlled Release Dosage Forms

Artificial skin: Artificial skin transplant is particularly needed in severely burnt cases such as: third degree

burns where patients lose more than 50 per cent of their skin layers, very large wounds and venous ulceration cases.

In theses situations, it is not possible for the skin to heal itself as a normal biological process due to the relatively slow nature of epithelial migration. Hence, the use of artificial skin transplant becomes mandatory.

Though using patient's own skin for coverage is very costly, yet highly effective in properly prepared wounds, it has drawbacks of hospitalization, immobilization, anesthesia and pain. Also some areas in the body are not amenable to grafting.

The best solution given by modern science of biotechnology as an effective alternative to auto grafting (using patient's skin or donors skin) is the artificial skin transplant. What is an artificial skin Artificial skin is a porous polymer film which is placed over an area of damaged skin (e.g. burnt skin) and host new skin. The edges of these polymers meet with healthy skin. Natural polymer collagen and synthetic polymer mesh.

Yet each one of these two techniques in its unique way has some disadvantages. Such as collagen not being strong enough to support new cells, whereas cells do not grow easily on synthetic mesh. To pool the benefits, these two materials are now combined to form hybrids that will be used in the artificial skin. There are two types of artificial skin: Xenograft - from non human donors, e.g. pig skin Allograft - from human donors.

Though xenograft can provide temporary coverage, it is always rejected due to immunological reaction; whereas, to avoid this immunological rejection allograft with genetic matching is always preferred. The general objective of this project was that to develop and clinically test new biomaterials, able to support the proliferation of dermal and epidermal cells, for use as composite scaffolds for the production of artificial skin .

For a better understanding of the objective it should be kept in mind that when considering tissue engineering in wound repair, the possible approaches generally fall into the following three categories:

Epidermal replacements -consisting of keratinocytes grown either alone (on the surface of a tissue culture flask), or in close association with a carrier vehicle such as a polymeric film or bioresorbable matrix.

Dermal replacements -consisting of a support structure able to support the infiltration, adherence, proliferation and neo-matrix production by fibroblasts (and in some cases endothelial cells).

Skin substitutes -are a combination of the above, able to support both dermal and epidermal components.

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Moreover, formulation of a suitable scaffold, able to facilitate the several biological processes needed for skin reconstruction is a major issue for groups developing dermal and epidermal substitutes. Suitability of the materials forming the scaffold is dependent upon several factors, these include the ability to:

support cell ingrowth provide a suitable substrate for adherence facilitate cell proliferation and production of extracellular matrix resorb from the wound site in a controlled fashion. further requirements include minimal toxicity, low immunogenicity and tensile properties

similar to uninjured tissue. composition and performance of skin substitutes must be evaluated before initiation of clinical

studies. Similar to other experimental therapies, cultured skin substitutes may be validated for function by evaluation in vitro and in animals.

Factors to be evaluated include: regulation of cellular proliferation and attachment, morphogenesis of cells into skin analogues, biochemical and biophysical assessment and histogenesis of skin analogues into functional skin. Based on the above considerations, the philosophy of the project was to develop materials and tissue engineering technologies to produce a complete skin equivalent consisting of both a dermal and an epidermal layer, an entirely autologous product, designed for the treatment of acute wounds (consequent to burn, trauma and surgery) and chronic wounds (diabetic foot ulcers, venous leg ulcers). The challenge was to develop tissue engineered materials which could be reproducibly developed at an industrial scale and could easily fit in the clinical practice of the end users. The idea is to combine fibroblasts (dermal cells) and keratinocytes (epidermal cells) in a scaffold made from a hyaluronic acid, a molecule naturally present in skin. The device would consist of a thin top membrane (for the keratinocytes) and an underlying sponge-like layer (for the fibroblasts). Designed to favour growth and joining of the two skin layers, the scaffold would be biodegraded in time with the healing process Scope of artificial skin

To develop new biocompatible materials based on biodegradable polymers which can be processed into tri-dimensional scaffolds.

To design appropriate configurations of the architecture of the scaffolds in order to enable effective cellular infiltration and proliferation

To determine the relevant mechanical, chemical, surface and biological properties of such materials in order to fulfill the essential requirements described in the EN 93/42 directive.

To test in vivo take rate and healing properties of the artificial skin grafts on animal models.

To carry out preliminary controlled clinical trials on burns and on scar/tattoo revision. To scale-up the technologies developed. To establish a more widespread culture on the clinical use of skin grafts in the treatment

not only of burn wounds and other acute wounds but also in the case of chronic ulcers. Concerning the originality of the project, important aspects are the choice to use a naturally

derived material (hyaluronic acid) for building the scaffold and to use autologous cells. The material for the scaffold:

To date, the vast majority of biomaterials are designed on the basis that implantable medical devices should be constructed of inert materials and their functional performance was restricted to simple mechanical or physical parameter. Recently, the conceptual approach to medical devices is changing and bio-interactivity is now the major requirement, rather than inertness. FAB:

FAB has dedicated much attention to the chemical derivatization of Hyaluronic acid (HA, Hyaluronan), with the aim of developing biopolymers which retain the biocompatibility of the parent molecule but have different physical properties (i.e. elasticity, viscosity and plasticity).

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Hyaluronan is a naturally occurring glycosaminoglycan, consisting of a repeating dimer of glucuronic acid and N-acetyl-glucosamine.

HA is a widely distributed and highly conserved polysaccharide, which plays an important role in a variety of vertebrate soft tissues, due to its peculiar physico-chemical and biological properties. By nature of its propensity to form highly hydrated and viscous matrices, HA imparts stiffness, resilience and lubricious faculty to various tissues.

The unique biophysical properties of HA are manifested in its mechanical function in the synovial fluid, the vitreous humor of the eye and the ability of connective tissues to resist compressive forces. HA is a major constituent of the extracellular matrix, where it has a profound influence on a variety of cellular events including cell migration and proliferation.

Its biological properties make HA an ideal candidate for the development of innovative biomaterials for various clinical indications. Purified medical-grade HA is currently employed to reduce the incidence of post-operative adhesions, as a viscoelastic aid in intraocular surgery, as a synovial replacement device, as a component of wound healing formulations and in various cosmetic preparations.

Numerous other potential applications are in reality precluded because of the physical nature of HA itself. Unmodified HA exists only in the form of aqueous gel which has a short residence time (i.e. it is rapidly degraded upon application) and is not susceptible to being manufactured into stable three-dimensional configurations.

FAB has optimized a number of well defined patented chemical modifications of HA that have been successfully employed to produce novel range of viscoelastic biopolymers These HA-derivatives can range from soluble material to solid hydrogel.

As the parent naturally occurring molecule, these HA-derivatives have demonstrated to offer the advantages of being recognized by cell receptors, of interacting with other extracellular matrix molecules and of being metabolized by intrinsic cellular pathways.

Moreover, these biomaterials have prolonged and controllable residence time upon application compared to the natural HA.

Being highly biocompatible, biointeractive, biodegradable and able to induce angiogenesis these materials have demonstrated to be among the most suitable candidates for the development of innovative 3-D scaffolds tissue engineering application. The cells source: The choice to use autologous cells as cells source for the artificial skin has the obvious benefit that there is little potential for rejection and the risk of disease transfer is minimal.

However, there is also the obvious need to scale-up and industrialize not just an of-the-shelf product but indeed a complete system/service for autologous skin tissue grafting. This means that there is the need to set up the logistic for carrying out effeiently the following step: -obtain a small biopsy from the patient and extract viable skin cells out of it -expand the cell number in certified high-tech laboratories -seed the cells on custom-made polymeric scaffolds -and deliver back the cell-biomaterial construct to be grafted onto the same patient.

Within the CE funded project we have indeed demonstrated the feasibility of this process. the production of artificial skin substitutes is a relatively new field, there are already some products and services available on the market.

Artificial skins are only a temporary fix; the patient will still need skin grafts.

However, with the use of artificial skins means a thinner skin graft, which allows the donor site and the patient to heal faster with fewer surgeries.

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The use of artificial skins has not yet been perfected and they are not right for every burn patient. Additionally scarring still results, but it may be less severe.

Dialysis membrane: Introduction Dialysis is Diffusion Through a Semi-Permeable Membrane

Dialysis is the movement of molecules by diffusion from high concentration to low concentration through a semi-permeable membrane.

Only those molecules that are small enough to fit through the membrane pores are able move through the membrane and reach equilibrium with the entire volume of solution in the system. Once equilibrium is reached, there is no further net movement of the substance because molecules will be moving through the pores into and out of the dialysis unit at the same rate. By contrast, large molecules that cannot pass through the membrane pores will remain on the same side of the membrane as they were when dialysis was initiated.

To remove additional unwanted substance, it is necessary to replace the dialysis buffer so that a new concentration gradient can be established.

Once the buffer is changed, movement of particles from high (inside the membrane) to low (outside the membrane) concentration will resume until equilibrium is once again reached.

With each change of dialysis buffer, substances inside the membrane are further purified by a factor equal to the volume difference of the two compartments.

For example, if one is dialyzing 1 ml of sample against 200 ml of dialysis buffer, the

concentration of the dialyzable substance at equilibrium will be diluted 200 less than at the start. Each new exchange against 200 ml of new dialysis buffer will dilute the sample 200 times more.

For example, for three exchanges of 200 ml, the sample will be diluted 200 x 200 x 200 or 8,000,000 times, assuming complete equilibrium was reached each time before the dialysis buffer was changed. Factors Affecting dialysis Rate

Factors that affect the completeness of dialysis include (1) dialysis buffer volume, (2) buffer composition, (3) the number of buffer changes, (4) time, (5) temperature and (6) particle size vs. pore size. Substances that are very much smaller than the pore size will reach equilibrium faster than substances that are only slightly smaller than the pores.

Dialysis Products use membranes that contain convoluted pores, not the tube-like pores often found in traditional dialysis tubing.

Convoluted pores allow small molecules to pass through the twists and turns of the pores with greater ease than do larger molecules. The greater the difference in molecular weight of the unwanted molecule vs. the molecular weight cut-off (MWCO) of the pore size, the greater the rate of dialysis. Dialysis Products are available in 3.5K, 7K and 10K MWCO membranes types.

The MWCO describes the molecular weight at which a compound will be 90% retained following overnight (17-hour) dialysis.

The MWCO is determined by testing many different proteins of known molecular weight. In general, these MWCO apply to globular molecules, such as most proteins. More linear proteins may be able to pass through the pores, even though their molecular weight exceeds the stated MWCO. To compensate for this, choose a dialysis device with a smaller MWCO.

For DNA or RNA, a MWCO no greater than one-third the MW should be used in order to prevent excessive sample loss.

The aforementioned factors highlight the difficulty in predicting dialysis time required for a particular sample. For thorough exchange of buffering salts, most researchers use a volume of dialysis buffer 200-500 times that of the sample and dialyze at room temperature for a total of 6-8 hours with 3 changes of the dialysis buffer.

After the last change, many researchers continue to dialyze overnight at 4°C. This general guideline is also appropriate for removal of most excess labeling reagents and

reaction byproducts from biotinylation or cross-linking experiments. However, some dyes form large aggregates and so do not dialyze efficiently despite having

low monomer molecular weights.

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Other Factors in Dialysis Because dialysis involves molecules passing through the semi-permeable membrane in

both directions and each substance reaches its own equilibrium independent of other substances, sample dilution can occur.

If a substance's concentration is higher on the outside of the membrane and is small enough to pass through the pores, it will have a net movement from the Dialysis buffer into the sample. Water is such a small molecule that it is capable of passing through the pores of virtually all dialysis membranes.

When dialyzing a high solute concentration against a dilute dialysis buffer, there will be a net movement of water (and possibly salts) into the dialysis unit through the membrane.

Glycerol and some sugars are especially hygroscopic, and as rapidly as they diffuse across the membrane to reach equilibrium, they also significantly affect the osmosis of water across the membrane and so may cause a change in volume of the sample.

Take care when dialyzing with large differences in glycerol or sugar concentration between sample and dialysis membrane. Prevent this movement of water and consequent change in sample volume by dialyzing in a "stepwise" fashion, minimizing the difference in water concentration between sample and dialysis buffer at each stage in the dialysis process.

For example, when processing a sample with very high solute concentration against a buffer with very low solute concentration, dialyze first against a fairly concentrated dialysis buffer. With each subsequent replacement of dialysis buffer, use a less concentrated buffer until the desired final buffer concentration is reached.

Finally, some molecules may stick to the dialysis membrane, which is made of regenerated cellulose. This nonspecific binding can result in sample loss.

The percent of total protein lost is partially dependent on the protein concentration. Protein loss caused by nonspecific binding to the membrane is negligible for concentrated

samples (>0.5 mg/ml) but may be significant with dilute protein samples (<0.1 mg/ml). Adding a “carrier” protein such as BSA to dilute protein sample before dialysis will

prevent this loss. Cellophane:

Cellophane is a thin, transparent sheet made of processed cellulose. Cellulose fibers from wood, cotton or hemp are dissolved in alkali to make a solution

called viscose, which is then extruded through a slit into an acid bath to reconvert the viscose into cellulose.

A similar process, using a hole (a spinneret) instead of a slit, is used to make a fibre called rayon.

Cellophane's low permeability to air, grease and bacteria makes it useful for food packaging.

Cellulose film has since been manufactured continuously since the mid-1930s and is still used today.

As well as packaging a variety of food items, there are also industrial applications, such as a base for self-adhesive tapes like Sellotape and Scotch Tape, a semi-permeable membrane in certain types of battery, and as a release agent in the manufacture of fibreglass and rubber products.

Typically, however, the use of the word "cellophane" has been genericized, and is often used informally to refer to a wide variety of plastic film products, even those not made of cellulose.

Cellophane sales have dwindled since the 1960s, through use of alternative packaging options, and the fact that Viscose is becoming less common because of the polluting effects of carbon disulfide and other by-products of the process.

However the fact that cellophane is 100% biodegradable has meant it is returning in popularity as a food wrapping.[1][2]