Biomedical application of functional polymers

Reactive & Functional Polymers 39 (1999) 99–138 Review Biomedical application of functional polymers Joseph Jagur-Grodzinski* The Weizmann Institute of Science, Rehovot 76100, Israel Received 11 May 1998; accepted 27 May 1998 Keywords: Biomedical application; Functional polymers; Drug delivery Introduction The application of polymeric materials for medical purposes is growing very fast Polymers have found applications in such diverse biomedical fields as tissue engineering, implantation of medical devices and artificial organs, prostheses, ophthalmology, dentistry, bone repair, and many other medical fields Polymerbased delivery systems enable controlled slow release of drugs into the body They also make possible targeting of drugs into sites of inflammation or tumors Prodrugs with macromolecular carriers have been also used for such purposes The term ‘prodrug’ has been coined to describe a harmless molecule which undergoes a reaction inside the body to release the active drug Polymeric prodrugs are obtained by conjugating biocompatible polymeric molecules with appropriate drugs Such macromolecular conjugates accumulate positively in tumors, because the permeability of cell membranes of tumor cells is higher than that of normal cells Subsequently, body enzymes ‘in situ’ catalyze the release of the drug Peripheral toxicities *E-mail: cpjagur@weizmann.weizmann.ac.il of the chemotherapies may thus be significantly reduced The use of polymers in the immunostimulation and for immobilization of enzymes have recently been probed The application of synthetic polymers for gene therapy has also been investigated They may provide a safer way of gene delivery than use of viruses as vectors Polymeric materials have also extensively been used for biosensors, in testing devices, and for bio-regulation Polymeric material suitable for a biomedical application must be ‘biocompatible’, at least on its surface Strictly speaking many polymeric systems used for implantation of medical devices into the body are considered to be ‘biocompatible’, though after implantation they become isolated from the tissues of the body by collagenous encapsulation They are, therefore, actually rejected by the body However, they not induce any harmful effects, thanks to encapsulation of the biofilm generated on the surface Interactions with the body are thus hindered Functional groups, properly located on a polymer as well as its structure, are usually responsible for its biocompatibility and / or biodegradability and may impart on it either therapeutic or / and toxic characteristics For example, 1381-5148 / 99 / $ – see front matter  1999 Elsevier Science B.V All rights reserved PII: S1381-5148( 98 )00054-6 100 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 carboxylic groups induce therapeutic activity of many drugs Cell and protein binding reactions and growth may strongly be affected by functional groups of an implanted polymer Encapsulation of an implant is triggered by adsorption of various proteins on its surface and adhesion of cells from the adjacent tissues The nature of such biofilm may be strongly influenced by the surface properties of the polymeric material Cell and protein binding reactions and growth of the attached cells can be effectively manipulated by appropriate functionalization of the surface of an implant As indicated above, an implanted polymeric material may be considered to be ‘biocompatible’, if its insertion into the body does not provoke an adverse reaction A thrombus is formed very fast when polymers contact blood cells Materials with non-thrombogenic bloodcompatible surfaces must, therefore, be used in contact with the blood stream Truly biocompatible polymers, used for medical purposes, should be able to recognize and cooperate in harmony with bio-assemblies and living cells without any non-specific interactions Biological molecules and synthetic ligands, designed to fit cell surface receptors, which are able to induce specific healing pathways should be immobilized on the surface to induce such effects To achieve these goals, appropriate functional groups should be incorporated into surfaces, which should resist non-specific adsorption of proteins [1] Biocompatible polymers used in biomedical applications must often be also biodegradable, and harmful products should not be generated as a result of their biodegradation Biodegradable polymers (also called bioerodible or bioresorbable) may be of synthetic or natural origin Non-toxic alcohols, acids and other low molecular products, easily eliminated by the body, are formed as a result of hydrolysis ‘in vivo’ of such biocompatible polymers High molecular weight polymers with hydrolytically unstable crosslinks may be bioeroded as a result of the release of their crosslinked chains Finally, water insoluble polymers may be converted into water soluble as a result of ionization, protonation, or hydrolysis of side chains Such conversion does not significantly affect molecular weight, but may be responsible for bioerosion in topical applications Progress in the field of tissue engineering depends on development of novel biocompatible fully bioresorbable polymers and introduction of processing techniques, which will enable reproducible three-dimensional architecture on the macro- and nanometer scale [2] Several useful biocompatible polymers of microbial origin are being produced from natural sources by fermentation processes They are non-toxic and truly biodegradable [3] Biodegradation is usually catalyzed by enzymes and it may involve both hydrolysis and oxidation Aliphatic chains are more flexible than aromatic ones and can more easily fit into active sites of enzymes Hence, they are usually easier to biodegrade Crystallinity hinders polymer degradation Irregularities in chain morphology prevent crystallization and favor degradation Extensive research efforts have been made to design appropriately functionalized, truly biocompatible and biodegradable polymeric systems Various surface treatments leading to its functionalization have been described in the literature Low toxicity glycol methacrylate resin embedding has been used for the immunohistochemical assessment of tissue response, by identifying inflammatory cells generated as response to contact with biomaterials A recently proposed procedure, based on infiltration step at 708C and curing phase at 208C enables processing tissues with implants in situ, with retention of both the immunoreactivity and biochemical potency of proteins and good preservation of morphology [4] Biological fouling of surfaces, due to adsorption of proteins and / or adhesion of cells, which may be highly undesirable in some cases, may be prevented by grafting onto surface such polymers as poly(N-isopropyl acrylamide) and its derivatives, or by deposition of a thin layer J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 of poly(ethylene oxide) (PEO) or its derivatives [5–7] Results of some of the recently conducted investigations of these subjects are discussed in the present review Drug delivery The release of drugs, absorbed or encapsulated by polymers, involves their slow and controllable diffusion from / or through polymeric materials It represents a well established approach to drug delivery Production of the slow release (SR) drugs by the pharmaceutical industry is nowadays a matter of routine Drugs, covalently attached to biodegradable polymers or dispersed in the polymeric matrix of such macromolecules, may be released by erosion / degradation of the polymer Both mechanisms may sustain the release of therapeutic agent by some systems Therapeutic molecules, complexed by polymers, may be released from gels by diffusion Release of the basic growth factor may be cited as an example of such system In others, the release mechanism may simply involve desorption of the adsorbed active agent Recent research efforts have been concentrated on systems capable to target drugs towards sick organs or cells Systems, suitable for local delivery in the vicinity of locations to be treated, have also been investigated Such strategies seem to provide the most effective approach to drug therapy, and may represent one of the main trends of the future Targeting can be attained by attaching biomolecules, capable to recognize specific cells, to the surface of nanoparticles containing a therapeutic agent In the absence of such recognition sites on their surface, less specific targeting can be achieved by using relatively large particles or macromolecules as vehicles of drug delivery Such particles or macromolecules accumulate preferentially in tumor cells, which are more permeable than healthy ones The term ‘prodrug’ has been coined as designation of a therapeutic 101 agent chemically bound to a another molecule, which becomes active upon its release Such release may be triggered by cell enzymes or by enzymes in the cell vicinity Local delivery may also be quite effective It is achieved by insertion of biodegradable implants in the vicinity of locations to be treated, or by introduction of catheters containing capsules either filled or coated with the slowly released drugs Biodegradable implants and nanoparticles with appropriate drugs dispersed evenly in the polymeric matrix (monolithic dispersion) as well as nanoparticles with a therapeutic agent adsorbed at the surface or loaded into the core, have been formulated for such applications 2.1 Biodegradable and biocompatible nanoparticles Polymeric micelles often self-assemble when block copolymers are used for their preparation Micelles, based on the biocompatible copolymers of poly(ethylene oxide (PEO) with poly( Llactic acid) (PLA) or with poly(b-benzyl-L-aspartate) (PBLA), have been described in the literature [8,9] Synthesis of such nanospheres with functional groups on their surface is summarized in Fig Aldehyde groups on the surface of the PEO– PLA micelles may react with the lysine residues of cell’s proteins They may also be used for attachment of the amino-containing ligands The hydroxyl groups on the surface of the PEO– PBLA micelles can be further derivatized and conjugated with molecules capable to pilot the modified micelles to specific sites of a living organism Such nanospheres have been tested as vehicles for delivery of anti-inflammatory and anti tumor drugs [10,11] 80–150 nm in diameter nanoparticles of the biocompatible and biodegradable polyester copolymer PLG [poly(lactide-co-glycolide)] (cf Fig 2) have been prepared by the nanoprecipitation method (they have been precipitated with acetone from their oily colloidal nanodispersion in water) Thus formed particles of PLG were 102 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 Fig (a) Poly(ethylene oxide)-co-b-benzyl-L-aspartate (PEO–PBLA) and (b) poly(ethylene oxide)-co-L-lactide (PEO–PLLa) micelles with aldehyde groups on their surface J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 Fig Poly(lactide-co-glycolide) (PLG) coated with 5–10 nm thick layer of the poly(propylene oxide)–poly(ethylene oxide) (PPO– PEO) block copolymer or with the tetrafunctional (PEO–PPO) -N-CH CH -N-(PPO– PEO) [12,13] Such coats are bound to the core of the nanospheres by the hydrophobic interactions of the PPO chains, while PEO chains protrude into the surrounding medium and form a steric barrier, which hinders the adsorption of certain plasma proteins onto the surface of such particles On the other hand, the PEO coat enhances adsorption of certain other plasma components In consequence, the PEO-coated nanospheres are not recognized by macrophages as foreign bodies and are not attacked by them [14] Derivatives of the phosphazene polymers, developed by Allcock and his co-workers, proved to be also suitable for biomedical applications [15–17] Long-circulating in the blood, 100–120 nm in diameter, PEO-coated nanoparticles of the poly(organo phosphazenes containing amino acid, have been prepared PEO–polyphosphazene copolymer, or poloxamine908 (a tetrafuntional PEO copolymer) has been deposited on their surface [18,19] Chemical formulae of such polyphosphazene derivatives are shown in Fig Arterial infusion of the PLG nanoparticles seems also to be a promising candidate for treatment of restenosis, caused by anginoplasty treatments [20] Polyethylene glycol (PEG)-coated nanospheres made from PLA, PLG, or other biodegradable polymers such as e.g poly(´-caprolactone) (PCL), may be used for the intravenous drug delivery PEG and PEO denote essentially identical polymers The only difference between the respective notations is that the 103 terminal hydroxyls of PEG may be replaced by methoxy groups in PEO It has been pointed out that PEG coating of nanospheres provides protection against interaction with the blood components, which induce removal of the foreign particles from the blood It prolongs, therefore, their circulation in the blood stream In consequence, thus coated nanospheres may function as circulation depots of the administered drugs [21,22] Obvious therapeutical benefits can be achieved by slowly releasing drugs into plasma, and thus altering their concentration profiles About 200 nm in diameter PEGcoated nanospheres, in which PEG is chemically bound to the core have been prepared, in the presence of monomethoxy PEG, by ring opening polymerization (with stannous octoate as catalyst) of such monomers as ´-caprolactone, lactide, and / or glycolide [22] Ring opening polymerization of these monomers in the presence of such multi functional hydroxy acids as citric, music, or tartaric, to which several molecules of the monomethoxy monoamine of PEG (MPEG–NH ) have been attached, yields multiblock (PEG) n –(X) m copolymers PEG–PLA copolymer in which NH terminated methoxy PEG molecules have been attached to tartaric acid is shown schematically in Fig It has been demonstrated that morphology, degradation, and drug encapsulation behavior of copolymers containing PEG blocks strongly depends on their chemical composition and structure Studies of nanoparticles composed of the diblocks of the poly( DL-lactide-co-glycolide) with the methoxy terminated polyethylene glycol (PLG–PEG) or of the branched multiblocks PLA–(PEG) , in which three methoxy terminated PEG chains are attached through a citric acid residue, suggested that they have a corecorona structure in an aqueous medium The polyester blocks form the solid inner core The anchored to them, on the surface, polyetheral (PEG) chains form a corona The nanoparticles, prepared using equimolar amounts of the PLLA–PEG and the PDLA–PEG stereoisomers, are shaped as discs with PEG chains sticking 104 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 Fig Polyphosphazenes for medical applications out from their surface Their hydrophobic / hydrophilic content seems to be just right for applications in cancer therapy and in gene therapy Such nanospheres are prepared by dispersing the methylene chloride solution of the copolymer in water and allowing the solvent to evaporate By attaching biotin to its free hydroxyl groups and complexing it with avidin, cell specific delivery may be attained NMR studies of such systems [23] revealed that the flexibility and mobility of the thus attached PEG chains is similar to that of the unattached PEG molecules dissolved in water PLG microspheres, with the PEG–dextran conjugates (cf Fig 5) attached to their surface, have been recently investigated as another variant of the above described approach 400–600 nm diameter spheres have been prepared To the glycopyranose hydroxyls of the dextran units targeting moieties can be attached [24] Nanoparticle of the biodegradable acrylic polymer poly(methylidene malonate) (PMM 2.1.2), which has ester groups as side branches, may also be useful for drug targeting [25] PMM 2.1.2 nanoparticles (cf Fig 6) have been prepared by anionic polymerization of ethyl - ethoxy - carbonyl - methyleno - oxycarbonyl acrylate in the presence of 1% dextran (pH 5.5 in phosphate buffer: 0.066 M KH PO / Na HPO , at 258C, 24 h) [26] Glycolic acid and ethyl alcohol are formed as a result of its hydrolysis Under the experimental conditions J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 105 Fig Multiblock (PEG) n –(X) m copolymers Amino terminated methoxy polyethylene glycol molecules attached to tartaric acid with PLa side chains they have not been found cytotoxic Another degradation pathway may, however, lead to formation of small amounts of formaldehyde Main drawback of the earlier investigated alkyl acrylates is their considerable cytotoxicity [26] Hydrogels suitable for delivery of pharmaceutically active proteins and peptides have recently prepared by attaching HEMA to dextran trough an oligolactate spacer [27] It has been assumed that biocompatibility of such polymers is good, since their degradation products are lactate, dextran, and HEMA Fig PEG–dextran conjugates 2.2 Local drug delivery Fig Poly(methylidene malonate) Ester groups as side chains (PMM 2.1.2) Implants from biodegradable polymers, encapsulating appropriate drugs or integrated with them, are usually used for internal local drug delivery This strategy enables delivery of a high local level of a drug at low level of systemic exposure and toxicity Thus hemorrhage complications of systemic delivery of an antithrombotic agent may be avoided by delivering it locally [28–31] Biodegradable, controlled-release, systems are gradually eliminated by 106 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 biological processes and not leave any residual implant structures Previously discussed in the section on nanoparticles biodagradable copolymers, such as PLG and poly(phosphazenes), as well as PLA, PGA, PCL, poly(orthoesters) (POE), poly(hydroxybutyrate), poly(alkane anhydrides), oxidized cellulose, collagen, gelatin, and various synthetic hydrogels, are used as biodegradable implants Gelatin and collagen, which biodegrade into amino acids and are eliminated from the body within a few days, as well as poly(alkane anhydrides), which biodegrade into aliphatic diacids and are eliminated from the body within a few weeks, are useful for short-term release PLG copolymers, eliminated from the body within 6–12 months, are also useful for relatively short-time drug release The rate of their degradation depends on the manufacturing procedure and on the composition of the copolymer Device elimination has been found to be much faster for the 50:50 than for the 85:15 PLG [29] Tests conducted on animal models revealed that, in spite of the fact that these copolymers are classified as biocompatible, they irritate coronary arteries of tested animals and cause inflammations Even more extensive inflammatory response has been found to be evoked by such supposedly biocompatible polymers as poly(hydroxybutyrate–valerate), POE, and PCL [32] Another work on PLG did not reveal, however, that implants of PLG in the brain tissue of rats cause adverse effects [33] These investigators concluded that PLG is suitable for brain implants Such implants, in the form of thin flexible strips of the poly anhydride copolymers, loaded with 2% of heparin and coated with PLA, effectively inhibit thrombosis and stenosis [28–31] NMR and EPR studies of their laminates with poly(FAD–SA) [where, SA stands for sebacic acid and FAD for fatty (erucic) acid dimer (cf Fig 7)] revealed [34,35] that the 50% FAD/ 50% SA mm thick tablets are almost completely absorbed few days after implantation From such PLA-coated laminates [28] heparin was released at a constant rate for at least 20 days On the other hand, the same copolymer uncoated by PLA released the absorbed by it heparin for days only The possibility of using fast eroding polyanhydrides laminated with the slowly eroding poly( DL-lactic acid) for manufacturing implants capable to release drugs in a programmed fashion has recently been investigated [36–38] Gopferich used an analogous FAD copolymer (cf Fig 8) poly[1, - bis( p - carboxyphenoxy)propane - co sebacic acid] 20:80 Mw 70 000 (PCPP–SA), coated with PLA’s (Mw 1900 and 17 400) He found, that such implants may release one drug in two phases or two different drugs one after another Poly(FAD–SA) matrices have been also tested on animal models as suitable materials for slow release of the anti-malaria Fig Polyerucic-co-sebacic acid (FAD–SA) Fig Polyh1,3-bis( p-carboxyphenoxy)propane-co-SAj (PCCP–SA) J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 drugs [39] Test on animal models indicate that implants of PCPP–SA loaded with the antitumor drug, carboplatin, may be effective in treating gliomas [40] Poly(´-caprolactone) (PCL) has been tested as a vehicle for slow release of drugs at tumor reaction sites It was used as a surgical paste loaded with the antitumor-drug taxol [41] An 80:20 copolymer of PCL with DL-lactide: [poly(CL–DLLA)] (see Fig 9) has also been found to be a suitable material for slow release devices In vitro experiments with this polyester revealed that up to 80% of the incorporated drug were released by the Mw 180 000 copolymer during 120 days It has been found in these experiments that a relatively slow release, which was initially diffusion controlled, became much faster after 50 days when its constant rate of release became governed by the rate of erosion of the polymer For a lower molecular weight copolymer (Mw 120 000), this transition was much sharper After the first 40 days, during which rates of release of the drug by the two copolymers were identical, the remaining drug was released in 20 days [42] Biodegradable-biocompatible ceramers may provide appropriate support for the growth of cells They have been prepared by reacting tetraethoxysilane with PCL end-capped with hydroxyls at both ends [43–45] Acid catalyzed hydrolysis of the so formed bis(triethoxysilane) PCL, followed by crosslinking with Si(OH) , yields a three dimensional network of a hybrid organic–inorganic ceramer (cf Fig 10) Such resorbable bioactive ceramers may be useful for the repair of skeletal tissues Polydepsipeptides may, in some cases, be preferable to the previously described polyester type biocompatible-biodegradable polymers 107 Concentrations of an acid formed locally upon the degradation of polydepsipeptides are, of course, lower than those formed as a result of the degradation of polyesters Moreover, their biodegradation may be induced by the action of both esterase and protease, and bioactive molecules may be attached to side groups of the a-amino acids The preparation of the polydepsipeptides by a new method of acylation of an amino acid with a hydroxy acid, followed by cyclization and ring opening polymerization of the product, have recently been discussed [46] Sequential polydepsipeptides had previously been prepared by reacting a-amino acids with phosgen Thus formed N-carboxy anhydrides (NCA) of amino acids have been subsequently reacted with 2-nitrophenylsulphonyl chloride (NPS-Cl) After purification of the product, its reaction with hydroxy acid yielded NPS-didepsipeptide Its reaction with the NCA derivative of another -amino acid yielded an end protected tridepsipeptide (cf Fig 11) This step could be repeated to form tetra-and pentadepsipeptides Thus obtained end-protected oligodepsipeptide was treated with HCl to convert it into a hydrochloride and purified by recrystallization Purified product was dissolved in DMSO, treated with a small excess of triethylamine and condensed at room temperature to yield a polydepsipeptide [47] The sequential polydepsipeptide AGL (cf Fig 12) has been tested as a biodagradable matrix for drug release The LH-RH agonist was incorporated into AGL matrix Its release from such implants was probed for treatment of prostates (on animal models) ‘In vivo’ experiments showed that such implants were completely biodegraded after 24 weeks 2.3 Inert biocompatible polymers as implants Fig Poly(caprolactone-co-DL-lactide) (P[CL–DLLa]) Implants with scaffolds made from polymeric materials, which are not biodegradable, have also been described in the literature: EVA (copolymer of polyethylene with poly(vinyl acetate), has been used for slow release of 108 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 Fig 10 Three dimensional network of a hybrid organic–inorganic ceramer growth-factors, antibodies, antigens, and other drugs [48] The asymmetric polyamide 4,6 membrane, encapsulating Langerhans islets, has been tested as a component of bioartificial pancreas [49] ‘In vivo’ and ‘in vitro’ tests suggested that such membranes, obtained by casting into water 15% solution of the polyamide in formic acid, have been biocompatible Permeability characteristics of such membranes have been improved by addition of 1% of PVP to the casting solution Polysulphone and poly(hydroxybutyrate) membranes, coated to increase their hemocompatibility by plasma deposition of a smooth fluorocarbon layer, have also been used for the bioartificial pancreas devices [7] Acrylic polymer Eudeagit L-55, which is pHsensitive (it dissolves in water at pH 5.5), has been used for the encapsulation of drugs to be delivered in the first part of the intestine [50] 2.4 Prodrugs with polymeric carriers It has been pointed out in the introduction to this review, that conjugation of drugs with soluble polymeric carriers may induce their accumulation in cancerous cells Tethers at which drugs are attached to macromolecules 124 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 be introduced into the body if their surface is thrombogenic It is generally believed that thrombus formation is initiated by adsorption of proteins from plasma followed by platelets adhesion and activation It seems that antithrombogenic characteristics may be induced by hydrophilic / hydrophobic micro domains on the surface Synthesis of the, phase separated hydrophilic / hydrophobic block copolymers, yields such materials In the bulk or on the surface of such block or comb like copolymers phase separation occurs spontaneously Several polyurethanes, known under the commercial names of Pellethanes and Biomers, are available for such medical applications For example, Pellethane 2363-80A is a poly(urethane-ether) elastomer (MDI / 1,4butanediol) with poly(tetramethylene oxide) (PTMO) forming soft hydrophobic segments Non-polar hydrophobic poly(isobutylene oxide) (PIBO) can also be used for the soft segments, instead of PTMO Good hemocompatibility was obtained when the content of (PIBO) was in the vicinity of 25% [105] Biomers are segmented poly(urethane-urea) elastomers (MDI / ethylenediamine) or (MDI / 4,49-diaminodiphenylmethane) with PTMO as soft segments Some grades may also contain other soft hydrophobic segments Biomer has been used as basic material for fabrication of such medical devices as artificial hearts and vascular grafts It is generally processed by repeated deposition of layers by solution casting The precise nature of the processing route may influence surface properties and mechanical response of the finished product and its in vivo response [106] Hemocompatibility of these phase separated polyurethanes may be further improved by surface treatments Radiation grafting, plasma discharge, and chemical treatments have been used as means for surface modifications [65] It has been mentioned in the introduction to this review, that non-thrombogenic characteristics of polymeric objects can be enhanced by grafting Nalkyl acrylamides onto their surface, by coating it or grafting onto it PEO, which is relatively unreactive towards plasma proteins and cells (cf refs [5–7]) Surface hemocompatibility can also be improved through grafting heparin (heparin is a polysaccharide which contains D-glucosamine and partially O-sulfated (at or C positions) D-glucuronic acid units), onto the surface of polymeric materials or by attachment of molecules resembling heparin segments [107] The surface modification approach have also been used to further improve the hemocompatibility of polyurethane block copolymers Significant reduction of thrombus formation was observed when a Biomer was coated by its blend with poly(N-isopropyl acrylamide) (IPAm) and swollen with heparin [108] IPAm is thermoresponsive (see previous section on hydrogels), and such blends release heparin at body temperature The grafting of acrylic acid onto poly(etherurethane-urea), synthesized from (1.3:1:1) MDI: poly(tetramethylene glycol) (Mw 1000): ethylene diamine as chain extender, activated by treatment with oxygen plasma, enables subsequent attachment of heparin (cf Fig 28) As a result of such treatment, activation of plasma proteins and platelets has been effectively suppressed and material with good in vitro blood compatibility has been obtained [109] The hemocompatibility of polyurethane has been improved by photo immobilization of dipyramidole, a powerful inhibitor of platelet activation and aggregation [110] Three of the four hydroxylic groups of dipyramidole were protected by tert-butyl-dimethylsilyl (TBDMS) groups and the photoreactive 4-azidobenzyl group was attached to the forth hydroxyl through a short triethylene glycol hydrophilic spacer Thus formed dipyramidole containing molecule has been immobilized on the surface of polyurethane (Pellethane D55) by short irradiation with UV-light (cf Fig 29) Clotting time has been substantially prolonged when treated Pellethane was exposed to blood plasma rich in platelets It had been noted several years ago that graft J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 125 Fig 28 Grafting of heparin onto the surface of poly(ether-urethane-urea) of poly(ethylene oxide) (PEO) onto surface of Biomer improves its blood compatibility [111] More recently it was observed that adsorption of fibrinogen and of g-globulin decreased with increasing grafting levels of N,N-dimethylacrylamide (DMAA) onto surfaces of poly(tetrafluoroethylene) (PTFE) and of poly(ethylene) (PE) It was suggested that such grafts may be suitable for long-term use in cardiovascular systems [112] Poly(dimethylsiloxane) (PDMS)–aramid multi-block copolymer (PAS) shows good blood compatibility because of its microphase separation and enrichment of PDMS segments in the outermost surface Tensile strength of the aramid is reduced drastically by the its copolymerization with PDMS Novel disiloxane– aramid multiblock copolymer with short dipropylamino-tetramethylsiloxane units (DATS) (cf Fig 30) combines, at 26 wt% of DATS in its bulk, the high tensile strength of an aromatic polyamide with the blood compatibility of PAS [113] Platelets scarcely adhere to the surface of such copolymer, because its outermost surface is enriched by DATS segments, which have a very low surface free-energy Biodegradable stents made from poly( DL-lactide-co-glycolide) (PLG) became compatible with blood when heparin had been immobilized on their surface [114] Immoblization was achieved by incubating PLG for h in 6% aq solution of heparin 3% of glutaraldehyde 10 Novel biodegradable polymers and improvements in older systems The term biodegradable polymers is used to describe polymers which are degraded either in the human body or by bacterial attack in the environment However, environmentally biodegradable polymers are not the subject of the present review Biodegradable polymers suitable for medical applications must be biocompatible Their degradation in the body is usually due to an enzymatic hydrolysis Drug delivering systems, vascular prostethis and orthopedic surgery are main fields of their clinical applications Medicinal components can be attached to a biodegradable polymer enabling prolongation of their therapeutic effect Heterogeneous biodegradation of a polymeric matrix, from its surface towards inner layers, is 126 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 Fig 29 Grafting of dipyramidole, through PEO tether, onto the surface of Pellethane D55 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 127 Fig 30 Siloxane–aramid copolymer called bioerosion Such degradation mode is preferable for drug delivery, as it provides a way for the release of drugs in the body at a constant rate with zero-order kinetics PLA, its homologous polymer PGA, and their copolymers PLG, are the most extensively used, clinically established biodegradable polymers PLA being preferred over PGA when slower degradation rates are required Extensive research efforts are directed towards developments of new or modified materials with further improved or, to specific applications adapted, biodegradation characteristics and mechanical properties Of great interest are also materials bearing pendant functional groups to which, molecules providing the desired hydrophilicity as well as drugs regulating cell functions, can be attached A process for production of high strength poly( L-lactic acid) (PLLA) fibers has been recently developed by combining melt extrusion with hot drawing [115] PLLA fibers with tensile strength as high as 0.87 GPa, yield strength of 0.29 GPa, and tensile modulus of 9.2 GPa, were obtained by drawing 10 times fibers collected at m / with initially low crystallinity As a result of hot drawing at overall draw ratio of up to l 190, their crystallinity increased to 65% Bioerodible, biocompatible polymers, which show zero-order kinetics for release of drugs, have recently been synthesized by bulk copolymerization of DL-lactide with adipic anhydride (ADAN) with stannous octoate as catalyst (48 h at 1608C)] ADAN-co-LA copolymers may be suitable for formulation of peptide and protein drugs Hydrophobic copolymers obtained at higher ADAN content exhibit zero-order kinetics of drug release [116] Polymers based on polyanhydrides seem to be of particular interest because they not induce inflammatory reactions, and the by-products formed during their biodegradation are non cytotoxic and nonmutagenic Moreover, as indicated above, some hydrophobic polyanhydrides display zero-order kinetics of drug release When aliphatic chains are replaced by aromatic rings, the biodegradation rates become much slower Living ring-opening ‘coordination-insertion’ polymerization of adipic anhydride (ADAN), initiated by aluminum triisopropoxide [Al(Oisopropyl) ] plus equivalent amount of a Lewis 128 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 base (pyridine or nicotine), has recently been described in the literature [117] It proceeds by insertion of the monomer into the ‘metal-alkoxide’ bond, while the ‘acyl-oxygen’ bond of ADAN is selectively cleaved The ionic character of the Al–O bond increases when a Lewis base is added to the Al-alkoxide catalyst Reasonably high molecular weight ADAN polymers (Mn 38 000), with rather narrow molecular weight distribution (Mw / Mn 1.25), were obtained by melt homopolymerization in the bulk of the ADAN (Over 99% yield after 10 at 808C and nicotine as a co-catalyst) (cf Fig 31) The rate of polymerization increases by order of magnitude as a result of addition of nicotine At the same time, transacetylation reactions which depend on coordination of carbonyl groups to Al atoms, are delayed In consequence, polymerization may be completed very fast without interference from transacetylation To avoid broadening of the molecular weight distribution, the reaction mixture must, how- Fig 31 Ring opening coordination-insertion polymerization catalyzed by the Al-alkoxide Lewis base complex ever, be de-activated immediately after completion of the polymerization Copolymers of ADAN with ´-caprolactone can be synthesized (cf Fig 32) using the sequential polymerization approach ‘Living’ PCL polymer, formed at 08C, upon addition of Al(O-i.Pr) to the solution of ´-caprolactone in toluene, initiates the polymerization of adipic anhydride, which may be conducted at 208C in the same solution Diblock copolymers with Mw / Mn 1.2 have thus been obtained Molecular weight of the PCL block could be as low as 1000 and of the PADAN block as high as 38 000 [Al(O-isopropyl) ] has also been used to initiate ring-opening polymerization of 1,4,8trioxa [4,6]spiro-9-undecanone (TOSU) as well as copolymerization of ´-caprolactone (´-CL) with lactic acid or TOSU [118,119] Such ‘living’ polymerization of TOSU is first order in monomer and catalyst concentrations Its copolymerization with ´-caprolactone initiated by the ‘living’ PCL chains yields narrow molecular weight block copolymers (Mw / Mn 1.2– 1.25) Deacetylation of the spiro group followed by subsequent reduction with NaBH , yields hydroxylated polymers (cf Fig 33) It has been suggested by Jerome that the presence of the reactive pendant hydroxyls on biodegradable and biocompatible polymers has great potential for applications in medicine, surgery and tissue engineering, and paves way to novel molecular architecture He proposed to use such hydroxyls for the synthesis of comb- Fig 32 Block copolymer of adipic anhydride with ´-caprolactone (PADAN–PCL) J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 129 Fig 33 Synthesis of functionalized polylactones like graft copolymers of lactones and lactides obtained by ring opening polymerization initiated by Al(C H ) in excess He also suggested that an additional grafting reaction may provide a route for formation of dendrimers Biodegradable poly(ester amide)s have been prepared by reacting 1,6-hexanediamine with di-methoxytartaric acid and the diacid formed by esterifying one mol of 1,6-hexanediol with two moles of succinic anhydride [120] Hydroxylic groups of those acids were protected by the pentachlorophenyl groups, while the trimethylsilyl groups were attached at the two ends of the diamine The polycondensation reaction (actually transamidation) was carried out at room temperature in chloroform The resulting polymer (cf Fig 34) could be degraded under very mild conditions The molecular weight distribution of the thus synthesized polymers was, however, broad It is believed that it may be advantageous to find ways to narrow its polydispersity Susceptibility of these polymers towards degradation under mild conditions is due to the methoxy side groups, which increase the hydrophilicity and favor nucleophilic attack on carbonyls Studies of the degradation products by NMR seem to indicate that cleavage of ester bonds represents the main pathway of the degradation 11 Effect of functional groups and of steric factors on the biocompatibility and biodegradability The study of the effect of various functional groups on the biological activity has been investigated with the help of self-assembled monolayers For example, self-assembled on gold alkanethiolates: X(CH ) –SH, where X C(O)OH, C(O)OCH , CH OH, or CH , were used to compare effect of these groups on the growth of endothelial cell growth and protein adsorption [121] Cell growth increased in the following order CH OH , C(O)OCH , CH < C(O)OH It was, however, much faster Fig 34 Poly(esteramide)s obtained by polycondensation of a conjugate of hexanediamine–dimethoxytartaric amide with hexanediol– succinic acid ester 130 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 on tissue culture polystyrene than on anolyte, indicating that multiple functionalities may have a pronounced synergistic effect Analogous approach has been used to study effect of various functional groups on apatite formation induced by contact with blood [122] In this case alkanethiols terminated with H PO , C(O)OH, C(O)NH , OH, NH , and CH , have been used Results of these studies indicated | OH | that: H PO C(O)OH C(O)NH | | NH OH < CH It was concluded that apatite formation may be initiated by adsorption of calcium induced by complexation, in the surface layer, with negatively charged ions Hydroxylic groups introduced into the surface of the poly(aryl ether ether ketone) (PEEK) by the heterogeneously conducted reduction with NaBH in DMSO at 1208C, moderately improved the biocompatibility of this polymer [123] Hydroxyls could be further activated by reaction with p-nitrophenyl chloroformate Compounds containing an amine group (e.g., lysine) were attached to the activated surface through carbamate linkage (cf Fig 35) Reactive groups, such as OH, COOH, and NR , can be easily inserted onto poly(amidoamine)s (PAA’s), a family of polymers obtained by stepwise polyaddition of amines to bisacrylamides [124–126] They are formed under very mild conditions, usually in aqueous solutions, at room temperature, and not require addition of catalysts Primary monoamines as well as secondary diamines yield linear polymers, while cross-linked resins may be obtained by using primary diamines PAA– PEG block copolymers may be obtained by reacting them with polyethylene glycol Their grafts on human serum albumin may be used for the preparation of nanospheres for drug delivery The biocompatible water soluble PAA’s may be used as polymeric components of prodrugs Several PAA polymers form complexes with heparin A typical linear, relatively easily hydrolyzable PAA, obtained by reaction of 1,4-bisacryloylpiperazine with 2-methylpiperazine is shown in Fig 36 Quaternized tert-amino polymers obtained by reaction of bis-acrylic esters with secondary amines caped by thiols (cf Fig 37), act as powerful antimicrobial agents [124] Other tert-amino polymers structurally related to PAA polymers include poly(ester-amine)s, Fig 36 Polyamidoamine (PAA) formed by polycondensing 1,4bis-acryloylpiperazine with 2-methylpiperazine Fig 35 Functionalization and attachment of lysine to the surface of poly(aryl ether ether ketone) (PEEK) J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 131 Fig 37 tert-Amino PAA (1,4-bisethylthiopiperazine–bismethacrylic ester of ethylene glycol) poly(ester-thioether-amine)s, poly(sulfonethioether-amine)s, and poly(amido-phosphine)s The degradation of the quaternized PAA’s is quite fast, while quaternized poly(esterthioether-amine)s and poly(amido-thioetheramine)s, are quite stable End-functionalized, low molecular weight, amphiphilic, non-toxic, poly(acryloyl morpholine) and poly(N-vinyl pyrrolidone) (cf Fig 38), can be used as enzymes modifiers [124] They can be prepared by radical polymerization in the presence of such chain transfer agents as 2-mercaptoethanol, 2-mercaptoethylamine, or 2mercaptoacetic acid, which introduce terminal functional groups Grafting of catalase, tyrosin- ase, trypsin, and superoxide dismutase, by the carboxy terminated poly(acryloyl morphine) activated as succinimidyl ester, did not cause any loss of their enzymatic activity However, similarly grafted ribonuclease-A retain only 30% of its activity Water soluble graft of PVPOH onto superoxide dismutase retained 95% of the original enzymatic activity End-functionalized by hydroxyls PCL oligomers can be prepared by heterogeneous catalysis (Aluminum hydroxide dispersed on a porous silica) [127] Exchange reaction between grafted alcohols and free alcohol is induced by addition of alcohol in excess (cf Fig 39) Fig 38 End-functionalized poly(N-acryloylomorpholine)s and poly(N-vinyl pyrrolidone) 132 J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 Fig 39 Ring opening polymerization and grafting of the ´-caprolactone Competition between propagation and exchange Good biocompatibility and low immunogenicity of PLA has been discussed in the first part of this review However, PLA has no reactive side-chain groups and is water insoluble To modify its properties and introduce functional pendant groups, copolymers of lactic acid with glycine–( L)lysine (Glc–Lys) and glycyne–( L)aspartic acid (Glc–Asp) depsipeptides, have been prepared [128] Schematic description of the formation of the poly(LA– (Glc–Asp)] depsipeptide is shown in Fig 40 Poly(LA–(Glc–Lys)] depsipeptide has been synthesized in an analogous way Treatment for h at 10–158C with 25% HBr / acetic acid was required for its deprotection and release of free NH2 side groups Derivatization of the a,bpoly(N-hydroxyethyl)-DL-aspartamide (PHEA) and of the a,b-poly(asparthydrazide) (PAHy), Fig 40 Synthesis of the polydepsipeptide: poly[LA–(Glc–Asp)] J Jagur-Grodzinski / Reactive & Functional Polymers 39 (1999) 99 – 138 non toxic, biocompatible, and water soluble polymers, with glycidyl methacrylate introduces pendant double bonds which facilitate formation of networks [129] Relatively low g-radiation doses are sufficient for crosslinking functionalized PHEA and PAHy Thus obtained hydrogels seem to be suitable for sustained release of drugs [130] Enzymatic degradation of the synthetic poly(b-hydroxybutyric acid) (PHB) was found to be affected strongly by its stereocomposition and tacticity [131–133] Clear preference for the degradation of R-PHB was observed, while PHB with the S repeat units was essentially non degradable Higher rates of degradation of synthetic PHB with 67–77% of R content, as compared to the natural polyester which contains only units with R configuration, have been attributed to its reduced crystallinity Atactic or slightly isotactic racemic PHB was shown to be degraded by PHB polymerase more rapidly than the highly crystalline isotactic polymer of the same overall stereocomposition On the other hand, syndiotactic PHB was not appreciably degraded by PHB depolymerases In this case effect due to the unfavorable tacticity and moderately high crystallinity are responsible for the very slow degradation rates Preferential enzymatic degradation of isotactic PHB seems to be due to the preferential affinity of binding domains of the PHB depolymerases [132] Transfection ability of complexes formed by cationic polymers with DNA seem to depend strongly on their torroidal structure and compactness, determined by the electrostatics of interactions [134] Primary amines, protonated at neutral pH are present in all such polymers Ability to buffer pH between pH to pH seems to be the requirement for the high transfection levels Complexes of fractured dendrimers and of the polyethyleneimine, mediating high transfection, have minimal tendency to aggregate They form in solutions suspensions of small particles with diameters of 100 nm or less Fractured dendrimers and polyethyleneimine are highly branched in an asymmetric 133 fashion, with one or two arms extended from each branch point It has been suggested that flexibility of the polymeric chains may be required for high transfection [134] Molecular imprinting is a novel relatively simple technique, which can provide ‘tailor made’ binding sites for molecular recognition that mimic enzymes, antibodies or receptors Synthesis of such polymers, with biomimetic functions, involves mixing functional monomer with template molecules with which they form complexes A spontaneous preorganization around the template takes place Highly crosslinked networks, firmly supporting the preorganized binding site structure, are formed when cross-linking agents are added Template-fit pockets, in which the position of the binding sites and their allignement is optimally arranged for binding specific target molecules, are formed when template molecules are removed The molecular imprinting may be based on covalent as well as on the non-covalent interactions In the covalent systems, the nonomer-template complexes are formed by reversible covalent bonds, such as labile Schiff-base, esters etc Hydrogen bonding, electrostatic interactions, hydrophobic interactions, and metal coordination are responsible for complex formation in the non-covalent systems More details, about this promising technique based on the molecular self-assembling, can be found in an excellent review 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