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Shape-memory polymers as a technology platform for biomedical applications Expert Review of Medical Devices Downloaded from informahealthcare.com by 193.34.92.27 on 05/20/14 For personal use only.

Expert Rev. Med. Devices 7(3), 357–379 (2010)

Andreas Lendlein†1,2, Marc Behl1, Bernhard Hiebl1,2 and Christian Wischke1,2 Centre for Biomaterial Development, Institute of Polymer Research, GKSS Research Centre Geesthacht GmbH, Kantstr. 55, 14513 Teltow, Germany 2 Berlin Brandenburg Center for Regenerative Therapies (BCRT), GKSS Research Centre Geesthacht GmbH, Kantstr. 55, 14513 Teltow, Germany † Author for correspondence: Tel: +49 3328 352 450 Fax: +49 3328 352 452 [email protected] 1

Polymeric materials are clinically required for medical devices, as well as controlled drug delivery systems. Depending on the application, the polymer has to provide suitable functionalities, for example, mechanical functions or the capability to actively move, so that an implant can be inserted in a compact shape through key-hole incisions and unfold to its functional shape in the body. Shape-memory polymers, as described herein regarding their general principle, compositions and architectures, have developed to a technology platform that allows the tailored design of such multifunctionality. In this way, defined movements of implants triggered either directly or indirectly, tailored mechanical properties, capability for sterilization, biodegradability, biocompatibility and controlled drug release can be realized. This comprehensive review of the scientific and patent literature illustrates that this technology enables the development of novel medical devices that will be clinically evaluated in the near future. Keywords : biocompatibility • biofunctionality • drug-release • polymer • shape-memory • sterilization

Motivation for introduction of SMP to biomedical applications

Polymeric biomaterials are presently applied in implants, surgical instruments, extracorporal devices, wound covers and controlled drugdelivery systems. Each application requires a specific combination of material properties and functions. In many cases implants have to fulfil certain mechanical functions. Examples are hernia meshes or wound chips. When functions of tissues or organs are substituted by a polymer-based device, biofunctionality is required. The hemocompatibility of surfaces is essential for medical devices in contact with blood, such as artificial heart valves or tubes used in extra corporal devices. The realization of novel material functions enabled important innovations in medical engineering. Hydrolytic degradability was the key function for the creation of temporary implants, which support wound healing or regeneration of processes (e.g., scaffolds as temporary substitute of the extracellular matrix [ECM] [1]) as long as required. Finally, they degrade and their degradation products are either metabolized or excreted. Their diffusion characteristics or the erosion behavior of polymeric biomaterials were the prerequisites for the application of these substances as matrix www.expert-reviews.com

10.1586/ERD.10.8

materials in controlled drug-release systems. The rapid progress in the development of surgical techniques, especially in minimally invasive surgery, leads to more complex requirements for modern implants. Exemplary questions arising include: how can a bulky implant be placed in the body through a small key-hole incision; or how can an implant be adapted to the individual geometry of a patient? Here, the shape-memory effect (SME) as a novel functionality of polymers could potentially play an important role. Shape-memory polymers (SMPs) can change their shape in a predefined way on demand when exposed to a suitable stimulus. Here, we focus on thermally induced SMPs, which are triggered by heat. A bulky implant having a switching temperature (Tsw) between room temperature and body temperature could be inserted into the body through a small incision in a compressed or elongated temporary shape. As soon as the implant is placed in the body, it is heated to body temperature and changes into its bulky application-relevant shape. Furthermore, the temporary shape can be adjusted to the individual needs of the patient. Often the SME needs to be combined with other functions, such as degradability or drug release. Degradability

© 2010 Andreas Lendlein

ISSN 1743-4440

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necessitates a sufficient elastic deformability of the polymer network. The netpoints of the polymer network, which are connected by the chain segments, determine the permanent shape of the SMP and can be of chemical (covalent bonds) or physical (intermolecular interactions) nature. Consequently, we use the term ‘shape-memory polymer network’ as an umbrella term. While covalent cross-links can be obtained by suitable cross-linking chemistry, physical crosslinks require a morphology consisting of at least two segregated domains – for example, a crystalline and an amorphous phase. In such multiphase polymers, the domains related to the highest thermal transition temperature (Tperm) are called hard domains and are General principle of SMPs SMPs are polymer materials, which can be temporarily stabilized acting as physical netpoints. in a second, temporary shape by application of external stress and The required deformability for the SME can be reached by subsequent fixation [2–13] . This temporary shape is stable until an the chain segments, which must allow a certain orientation to be appropriate stimulus is applied to the shaped body, which induces obtained, which increases with growing length and flexibility of the recovery of the original shape (Figure 1A) . The movement dur- these chains. The recoiling of the chain segments, which is entroing recovery from the temporary to the original shape reverses the pically driven once the material is in the rubbery state, enables mechanical deformation, which was applied for the formation of the recovery of the permanent shape. As a consequence, the SME the temporary shape and is in this way predefined. requires the temporary fixation of the chain segments’ conformaThe SME results from a suitable molecular polymer network tion in the deformed shape for stabilizing the temporary shape. architecture in combination with a tailored processing and pro- Such a reversible fixation can be realized by the formation of addigramming technology. In Figure 1B a scheme for such a polymer tional reversible netpoints, which can be formed and cleaved on network, as well as the programming, is depicted. This program- demand. Like the permanent netpoints, these additional netpoints ming technology is also named as a ‘shape-memory creation process’ can be built by the formation of chemical (covalent) reversible (SMCP). Suitable polymer materials enabling SME require a poly- netpoints or by the solidification of the switching domains, which mer network architecture consisting of netpoints (dots in Figure 1B) are formed by the switching segments ( Figure 1B, middle). and molecular switches (lines in Figure 1B), which are sensitive to an Thermosensitive SMPs, which is presently the most promiexternal stimulus. The deformation creating the temporary shape nent class of SMPs, can be additionally classified according to the thermal transitions related to the solidification of the polymer chain segments. These polymer chain segments i and, when applicable, the associated domains, are acting as molecular switches due to their correlated thermal transition Extension °C temperatures (Ttrans ). These molecular & cooling switches must be able to fix the deformed Ttrans ii shape temporarily under conditions relevant for the particular application by Ttrans forming additional reversible cross-links. As already explained for the permanent netpoints, these additional temporary iii Ttrans Heating cross-links can be formed by physical interactions or by covalent bonds. Physical cross-linking is realized by solidification, which occurs on cooling, such as vitrification or crystallization. When the SMP is reheated, the crystallites will melt or the glassy polymer chain segments will Figure 1. Macroscopic effect and molecular mechanism of the thermally-induced shape-memory effect of a covalently cross-linked polymer with Ttrans = Tm. return to the viscous state. Therefore, in (A) Time series of photographs showing recovery of a shape-memory tube made of a thermoplastic SMPs, the domains associpoly(e-caprolactone)dimethacrylate polymer network (the Mn of the network’s switching ated to the polymer chain segments with 4 segments was 10 ) that had been programmed to form a flat helix. i–iii: Start to finish of the second highest Ttrans are acting as the process; total time, 10 s at 50°C. (B) Schematic illustration of a shape-memory switching domains. The reaction of two creation procedure and recovery process. Ttrans: Thermal transition temperature related to the switching phase. functional groups under formation of a Reproduced with permission from [2] . © Wiley-VCH Verlag GmbH & Co. KGaA. reversible chemical bond enables chemical

Expert Review of Medical Devices Downloaded from informahealthcare.com by 193.34.92.27 on 05/20/14 For personal use only.

would be a valuable additional contribution, as a second surgery for explantation could be avoided. In this article, the general principle of the thermally induced SME is described, and selected examples for biomaterials with SME are introduced. An overview of medical devices being developed based on SMPs is given, including general surgery, the cardiovascular system or the urogenital application. Special emphasis is given to the potential use of SMPs as matrix materials in controlled drug-release systems.

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Shape-memory polymers as a technology platform for biomedical applications

cross-links as switches. The application of a suitable stimulus cleaves this bond on demand, so that SMPs can be categorized by the type of molecular switching mechanism, accordingly. If these additional cross-links are based on physical interaction, a further distinction according to Ttrans can be made, which can be a glass transition Tg, a melting transition Tm or a liquid crystalline transition. While Tgs are extending over a broad temperature interval, the latter ones can show a transition in a relatively small temperature interval. Melting transitions mostly extend over a temperature range between 15 and 20 K, while liquid crystalline transitions can even occur in a temperature interval between 1 and 5 K. The shape-memory properties are quantified in cyclic, stimuli-specific mechanical tests that consist of the SMCP and the recovery of the original, permanent shape [2,14–16] . From the data obtained, the shape fixity ratio (R f ) and the shape recovery ratio (R r) can be determined. R f is a measure for the ability of the switching segment to fix a mechanical deformation – for example, an elongation to em, applied during SMCP, when the temporary shape is created. R r determines the ability of the material to memorize its permanent shape. Different test protocols have been developed. They differ in SMCP, which can be performed under constant strain or constant stress conditions. When the recovery process is performed under stress-free conditions, for thermally induced SMPs the switching temperature Tsw, triggering the macroscopic SME, can be determined. In addition, the temperature at which maximum stress is developed (Tsmax) can be determined from recovery processes under constant strain. Examples for shape-memory polymer systems intended for biomedical applications

With respect to SMPs, biomedical applications require certain properties or functionalities, which can be fulfilled by the appropriate choice of suitable SMP architectures. An example is a covalent SMP network in which the temporary shape is fixed only by one switching domain. Such SMP networks display a high decrease in the mechanical properties, such as Young’s modulus or strain at break, when the temperature is exceeding Ttrans and the SME is triggered. In contrast, thermoplastic SMPs, in which a second phase provides the physical cross-links determining the permanent shape, do not show such a drastic change in the mechanical properties, as the remaining crystalline domains reinforce the amorphous switching domains. Another important point to be considered from an application point of view is the processability of the different SMP architectures. While covalent SMP networks provide in many cases higher R f and R r compared with similar thermoplastic architectures, thermoplastic SMPs provide the advantage of processing by extrusion. Besides these aspects for the creation of SMP materials, the underlying polymer network architecture has a significant influence with respect to the application. When actuated, thermoplastic SMPs are able to develop much higher stresses compared with covalent SMPs due to the much higher Young’s moduli, which can be achieved during the SMCP. In conclusion, before application, it needs to be decided if higher values of R f and R r or higher stresses are more important. Additionally, SMPs can be considered as www.expert-reviews.com

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multifunctional materials when, for example, the SME and hydrolytic degradability are present in one material. Multifunctionality is the unexpected combination of material functions [17] . A last example, which needs to be considered for SMPs for biomedical applications, is the need of appropriate sterilization. This requires, for example, a Ttrans of the SMP, which is higher than the temperature attained during the sterilization technique. Furthermore, in case of sterilization techniques using reactive components, building blocks, which are nonreactive towards these components, have to be used. As a result, several SMP network architectures have been invented, each having certain advantages. Representative examples will be shown in this chapter, following a categorization by Ttrans [2,3,5,6] . Physically cross-linked shape-memory polymers Multiblock copolymers with Ttrans = Tg

An important group of physically cross-linked shape-memory polymers are linear block copolymers. As most polymer segments with a sufficient difference in the molecular structure are not mixable with each other, multiblock copolymers, in which Ttrans relies on a glass transition temperature (Ttrans = Tg), can be obtained. Thermoplastic SMPs can be synthesized by direct coupling of presynthesized polymer blocks with a reactive linker, by applying the prepolymer method or by melt blending. The application of the prepolymer method enabled the production of thermoplastic polyurethane elastomers on an industrial scale. In this process, isocyanate-terminated prepolymers are obtained by reaction of hydroxytelechelic oligo-esters or -ethers with an excess of a low-molecular-weight diisocyanate. Examples of such polymers are polyetherurethanes (PEUs) with a hard segment from methylenediisocyanate (MDI), and butanediol and poly(tetramethyleneoxide) (PTMO) or poly(ethylene adipate) as a second segment [2] . In most cases these polymers form a mixed domain acting as a switching phase. The quality of phase separation between the polyurethane segments and the PEU is determined by the molecular weight of the polytetrahydro furandiols used as educt. In a commercially available PEU synthesized from methylene bis(p-cyclohexyl isocyanate) (H12 MDI), 1,4-butanediol, and PTMO diol, Tg is at 74°C. This material is used in artificial hearts, wound dressings and pacemaker leads [18] . Determination of shape-memory properties revealed an R f of 100% and R r of 80% after the third cycle at em 50%. The incorporation of silica-coated magnetic nanoparticles of iron(III)oxide core into Tecoflex® enabled the remote actuation of the thermally induced SME in alternating magnetic fields [19] . Multiblock copolymers with Ttrans = Tm

Prominent examples for block copolymers with Ttrans = Tm are PEU, polyesterurethanes (PESU) and polyetheresterurethanes. In PESU, the oligourethane segments are acting as hard segments, while the polyester segments, for example, poly(e-caprolactone) (PCL), are forming the switching segments. Biocompatible and simultaneously biodegradable multiblock copolymers with shapememory properties could be synthesized via a co-condensation 359


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of diols from poly(p-dioxanone)diol and PCL segments using an isomeric mixture of 1,6-diisocyanato 2,2,4-trimethyl hexane and 1,6-diisocyanato-2,4,4-trimethylhexane (TMDI) as a bif unctional coupling agent [20] . The required diols were synthesized by ring-opening polymerization of the cyclic esters [21] . In these polymers, named PDC, the domains resulting from poly(p-dioxanone) (PPDO) segments determine the permanent shape, while the crystallizable PCL segment is acting as the switching segment. Hard segment contents of the synthesized polymers ranged from 0 to 83 wt%. The mechanical properties strongly depended on the hard segment content. The multiblock copolymers were elastic at room temperature and exhibited high elongation at break eR up to 1000%, but an increase in the amount of PPDO resulted in a stiffer polymer and a decrease of the corresponding eR. R f between 98 and 99.5% was determined throughout all cycles, while Rr depended on the cycle number and gradually approached 100%. In multiblock copolymers having PCL switching segments and polyhedral oligosilsesquioxane (POSS) moieties in the polyurethane hard segments, a tremendous increase in the elasticity above Ttrans was determined, which was attributed to physical crosslinks formed in the hard block through POSS crystallization [22] . Recently, the degradation characteristics of crystalline multiblock copolymers have been investigated on shape-memory polyurethane based on poly(adipate)diol (Mw 3500 gmol-1) as a switching segment and a hard segment derived from methylenebis(4-phenylisocyanate) (MDI) and butanediol. The degradation process could be divided into three phases: an induction phase, a phase of continuous degradation and a phase of accelerated degradation. Rr remained fairly constant during phase one and decreased slowly during phase two. The increase in crystallinity in phase two was accompanied by an increase in Rf [23] . During the hydrolytic degradation of polyester segments, carboxylic acids are generated. In contrast, when some of the ester functional groups are replaced by peptide segments as in polydepsipeptides, during degradation the carboxylic acid can be directly buffered by the amino function. The substitution of a polyester segment by a polydepsipeptide segment in multiblock copolymers is thought to combine advantageous degradation behavior of the depsipeptide segment with the shape memory capability of multiblock copolymers with a PCL switching segment. Thermoplastic multiblock copolymers with polydepsipeptide and PCL segments providing shape-memory capability were synthesized via coupling of oligodepsipeptide diol and PCL diol using TMDI [24] . In these polymer materials, the switching domains were formed by the PCL block, which is acting as switching segment, while the domains determining the permanent shape were formed by the polydepsipeptide segments. The shape-memory properties of such a thermoplastic multiblock copolymer with 50 wt% of PCL segments yielded an Rf and Rr of more than 96% for all cycles, with a Tsw around body temperature at em = 50%. Covalently cross-linked shape-memory polymers

Shape-memory polymers networks providing covalent netpoints can be obtained by cross-linking of linear or branched polymers, 360

as well as by (co)polymerization/poly(co)condensation of one or several monomers, whereby at least one has to be at least tri-functional. Depending on the synthesis strategy, cross-links can be created during synthesis or by post-processing methods. Cross-linking by post-processing methods is described in [2,6,7] . The shape-memory behavior of covalently cross-linked polymer networks can be controlled by certain general parameters, such as the functionality of the cross-links, the network chain segment length, the nature of the switching segment or the number of phases. The mechanical properties are controlled by the cross-link density, which is a function of the chain segment length and the functionality of cross-links. In contrast, the nature of the switching segment influences the hydrolytic degradation, as well as the characteristics of the shape-memory effect, which can extend either over a broader temperature interval (Tg) or a sharp melting transition (Tm). Ttrans = Tg

Covalently cross-linked polymer networks with Ttrans = Tg were obtained as AB copolymer networks from the copolymerization of diethylene glycol dimethacrylate (DEGMA) or poly(ethylene glycol) dimethacrylate with low-molecular-weight acrylates. In these polymers the elastic properties, as well as Ttrans, could be individually controlled by the choice of the acrylate. as well as by the molecular length and by the weight amount of the cross-linker [25–29] . In AB copolymer networks obtained from the copolymerization of DEGMA with t-butyl acrylate (tBA), the Tg was around 55°C. The variation of the cross-linker content between 0 and 40 wt% enabled control of the rubbery modulus, and could be adjusted between 1.5 and 11.5 MPa [26] . When methacrylate was used in such AB copolymer networks instead of t-butyl acrylate, the Tg could be varied between 56 and 92°C, while values of the rubbery modulus between 9.3 and 23.0 MPa were obtained [27] . Furthermore, Fe3O4 particles can be added to the acrylate mixtures, enabling remotely controllable SMP network composites [30] . Recently, the influence of the cross-link density or the chain flexibility on the Tg of AB copolymers was systematically investigated in two polymer systems based on the cross-linking reaction of oxiranes and amines. In the first system, the crosslink density was reduced, while in the second system the chain flexibility was increased. Both approaches allowed the variation of Tg between 6 and 89°C [31] . Amorphous, covalently cross-linked polymers can also be synthesized by polyaddition or polycondensation reactions [32,33] . Cross-linking of star-shaped hydroxytelechelic co-oligoesters yielded copolyesterurethane networks with Ttrans = Tg. Ringopening polymerization of a-hydroxy acids with hydroxytelechelic initiators introduced the cross-link points in co-oligoester segments, which were formed by copolymerization of diglycolide and rac-dilactide and yielded oligo[(rac-lactide)-co-glycolide] (oCG). 1,1,1-tris(hydroxymethyl)ethane or pentaerythrite were used as initiators. The use of 1,1,1-tris(hydroxymethyl)ethane resulted in trifunctional netpoints, while usage of pentaerythrite yielded tetrafunctional netpoints. The final polymer networks were obtained by addition TMDI to the co-oligoesters and subsequent heating. R f and Rr of all prepared networks were higher than 90% for all Expert Rev. Med. Devices 7(3), (2010)



cycles. em was varied between 50 and 100%. For the first cycle, values were mostly lower compared with the second cycle, which was attributed to a segment-chain orientation effect. Urethane bonds in such SMP networks could be avoided by direct condensation of a dicarbonic acid (dodecanedioic acid) with a polyol(glycerol) [33] . The substitution of the glycolide copolymer by other cyclic diesters in the synthesis of hydroxytelechelic co-oligoesters was shown to be another parameter to control the Ttrans of such amorphous polymer networks [34] . The mechanical properties of such polymer networks could be substantially enhanced by the introduction of an additional amorphous phase, which results from phase separation and, depending on the applied synthesis strategy, can be located either in the hydroxytelechelic co-oligoester or separately [35] . Transparent and hydrolytically degradable SMP networks with Ttrans = Tg based on acrylate chemistry could be obtained by UV polymerization of poly[(l-lactide)-ran-glycolide] dimetha crylates. The macrodimethacrylates are prepared via a two-step synthesis: hydroxy telechelic poly[(l-lactide)-ran-glycolide]s (PLG) with number average molecular weights (Mn) between 1000 and 5700 gmol-1 were prepared by ring-opening polymerization from l,l-dilactide, diglycolide and ethylene glycol as initiator, using dibutyltin oxide as the catalyst [36] , and afterwards functionalized with methacryloyl chloride resulting in terminal methacrylate groups. The application of solid state NMR using a 13C signal at 44 ppm enabled determination of the real cross-linking density, whereas the determination by the degree of swelling is limited because of constraints such as physical entanglements [37] . Tg was shown to be almost constant at approximately 55°C. In contrast, the mechanical properties at temperatures higher than Tg depended on cross-linking density [38] . Good shape-memory properties with R r close to 100% were obtained under stress-controlled programming. As already shown for SMPs with Ttrans = Tm, the elastic properties of such SMP networks could be significantly improved by the incorporation of a second amorphous phase, yielding AB networks [39] . Ethylacrylate, BA or hexylacrylate were chosen as monomers providing the second phase, which keeps the material elastic. Although phase separation could not be achieved in most cases, the limitations in the mechanical properties of networks based on DM could be eliminated in general. Ttrans of the shapememory effect could be varied between 9 and 45°C by the choice of the acrylate comonomer and the comonomer ratio. For the networks containing polyethylacrylate as second phase, R f and R r values were over 97 and 98.5%. For strengthening, the phase-separation properties of the second phase providing additional elasticity, SMP networks with an ABA architecture were designed. In these ABA polymer networks synthesized from poly[(rac-lactide)-block-polypropyleneglycolblock-(rac-lactide)]dimethacrylate precursors, a Tg of the phase provided by the poly(propylene oxide) of 50°C and a Tg of 50°C of the phase resulting from the poly(rac-lactide) was expected. In these polymer networks a distinct phase separation of the resulting polymer networks could be observed for macrodimethacrylate precursors with an Mn more than 10 kD [40] . An additional transition www.expert-reviews.com

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associated with the mixed phase between the phase transition resulting from the poly(propylene oxide) and the poly(rac-lactide), as well as the phase transition from the poly(propylene oxide), was observed, when Mn of the macrodiol-dimethacrylate precursors was less than 10 kD. Ttrans = Tm

Similar to amorphous covalently cross-linked polymer networks based on acrylates, polymer networks with Ttrans = Tm can be synthesized based on PCL diol as the component that forms a crystallizable switching segment [41] . For that purpose, PCL diols [36] have been functionalized with methacrylate end groups that can undergo a polymerization reaction. The cross-linking was performed without the addition of an initiator by photopolymerization so that harmful substances can be avoided. The number average molecular weights of the PCL dimethacrylates (PCLDMA) used varied between 1500 and 10,000 gmol-1. Ttrans of the materials was adjustable via the molecular weight of the macro dimethacrylates used in the synthesis, and was varied between 30 and 50 °C. The photoset materials, named N-PCLDMA, displayed excellent shape-memory properties, with R r between 92 and 97% and Rf between 86 and 97% after five cycles. In addition, PCLDMA could be used for the preparation of degradable SMP network composites and those SMEs can be triggered by a magnetic field. In this case, the SMP network matrix needs to be prepared by a thermal radical initiator, as the absorption of the particles prevents initiation [42] . The in vitro degradation of such SMP network nanocomposites consisting of cross-linked poly(e-caprolactone) and Fe3O4 nanoparticles has recently been studied [43] . Other covalently cross-linked SMPs with Ttrans = Tm based on semi-crystalline polycyclooctene could be obtained by ring-opening methathesis [44] . This polymer contained unsaturated carbon bonds, which could be cross-linked after processing by the addition of dicumyl peroxide. The loading of the SMP with thermochromic dyes enabled SMPs with built-in threshold sensors [45] . The hydrolytic degradability of polyesters can be substantially increased by the incorporation of additional easily cleavable ester bonds. Crystalline, hydrolytically degradable covalently cross-linked shape-memory polymer networks (N-PCGDMA) with an increased hydrolytical degradation were obtained by UV-curing of poly[(e-caprolactone)-co-glycolide]-dimetha crylates (PCGDMA) [46] . The precursors had a glycolide content of 14 mol% and an Mn of 4900 gmol-1 or 12,800 gmol-1. While Tsw of N-PCGDMA using PCGDMA of 4900 gmol-1 was at 30°C, N-PCGDMA built from PCGDMA of 12,800 gmol-1 had a Tsw at 38°C. N-PCGDMA using PCGDMA of 4900 gmol-1 displayed Rf greater than 96% and Rr of 86% after five cycles with em = 150%. With increasing molecular weight of the precursors, R r and R f increased, so that R f of greater than 97% and R f of greater than 94% were obtained. The elastic properties of covalently cross-linked polymer networks can be enhanced by the addition of an additional elastic phase provided by a second co-monomer [47] . The co-monomer of choice was n-butyl acrylate (BA) because of the low Tg value of 361


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pure poly(n-butyl acrylate) (-55°C). The molecular parameters for controlling the switching temperature and mechanical properties were the molecular weight of the used PCLDMA cross-linker and the comonomer content of BA. The number average molecular weights of the PCLDMA used were 2000 and 10,000 gmol-1. With decreasing content of butyl acrylate, the cross-link density increased. The butyl acrylate content influenced the thermal properties of the AB network formed, especially for the PCLDMA with a molecular weight of 2000. In this series, a Tm = 25°C could be observed only in the case of a content of 11 wt% of BA; the other networks of this series were found to be completely amorphous. In contrast, the homo-network of PCLDMA has Tm = 32°C. For the networks of BA and PCLDMA having a number average molecular weight of 10,000, the melting point decreased by up to 5 K to 46°C with increasing BA content. The mechanical properties of both materials containing PCL segments with number average molecular weights of 2000 or 10,000 changed significantly with increasing BA content. The values of the elastic modulus, the tensile strength em and the tensile stress at break sR decrease by approximately an order of magnitude. In contrast, the absolute values could be increased by one order of magnitude higher by increasing the number average molecular weights of the PCL segments from 2000 to 10,000. For polymer networks containing the higher molecular weight PCL segments, Rf between 95 and 85%, as well as Rr between 98 and 93%, was determined. Sterilization & in vitro cytotoxicity testing

The risk of infection caused by the implantation of medical devices has to be minimized, and biomaterials have to be nontoxic with contacting tissues, cells and body fluids (Council Directive 93/42/EEC). The international standard ISO 10993 (‘Biological Evaluation of Medical Devices’) provides guidelines for bio compatibility testings in vitro. Part five of the standard deals with ‘Tests for cytotoxicity – in vitro methods’. In vitro toxicity testing regimes attempt to measure cytotoxicity, cytocompatibility, cell adhesion, genotoxicity, hemocompatibility, cell proliferation and a host of other measurable cell parameters. In general, no single method can cover the complexity of general toxicity in humans [48] . In order to assess the potential toxicity of contaminants (catalysts, monomers, solvents and endotoxins), several in vitro tests have been developed to measure different biological end points [48–50] . The US Pharmacopeia metric (USP23-NF18) provides a helpful grading system for the cell test results. The actual accepted standard to minimize the risk of infection by medical devices made of SMPs is to make them sterile before use. Sterilization is defined as the act or process, physical or chemical, which destroys or eliminates all forms of life, especially microorganisms [51] . According to the European Standard EN 556-1 and the US FDA guide (2002) concerning the sterilization of biomaterials, after sterilization, a live microbe is allowed to be found in only one in 1 million samples of the sterilized load [52] . However, sterilization can generally alter the physical, mechanical and biological properties (biocompatibility, biofunctionality) of materials subjected to it [53] . The selection process of appropriate sterilization techniques for SMPs is challenging because of the variety of effects 362

(e.g., thermal transitions, transition temperatures and packing constraints) that can impact on SMP devices. Autoclave, or steam sterilization, is generally not used with thermoplastics because of its high temperature range (121–132°C), which can melt the polymer or alter its morphological structure [54] . This would result in an actuation of the SME during the sterilization process, as Ttrans is in general below this processing temperature required for the sterilization process. Furthermore, the high temperature can potentially damage drugs embedded within the polymer matrix. Among the methods listed in the US FDA guide, the low temperature sterilization (LTS) methods based on g- or e-beam-irradiation, hydrogen peroxide vapor, low-temperature gas plasma and ethylene oxide (EO) are used for polymeric materials. Among them, a careful selection has to be made, as hydrogen peroxide vapor or EO may react with certain chemical groups of the SMP. Possible examples are oxidation reactions, for example, between amine groups and hydrogen peroxide, or cross-linking reactions between EO and hydroxyl or amine groups. While these effects might be neglected for materials not intended to change their shape, in case of SMPs, such reactions might result in an undesired shift of Ttrans, either resulting in actuation of the SME during sterilization or preventing it at the application site, when stimulated by exposure to the intended temperature range. In addition, newly formed covalent bonds might also prevent the capability of material to display a SME. Irradiation-based sterilization techniques such as g- and e-beam sterilization are instantaneous, penetrating and nontoxic, but may be associated with changes in the molecular structure [55] . Irradiation, for instance, is known to reduce the glass temperature (Tg) and the molecular weight of PLG (80:20) and PCL (Mw = 80 kDa) [56–60] . The molecular weight of PLG, which is known as a candidate material for SMP synthesis, decreases after g-irradiation, and this degradation process continues on storage for 4 weeks [57,58,61] . In addition, e-beam sterilization (50–200 kGy) is described to alter the degradation properties of PLG [57,61] . For PCL, Filipczak et al. demonstrated by e-beam irradiation (9.0 kGy/min) that the molecular weight distribution becomes broadened through combined chain scission and cross-linking [56] . A fairly new LTS technique to rapidly sterilize shape-memorybased medical devices is the low-temperature plasma (LTP) sterilization. Strong surface changes were seen after LTP sterilization of polyurethane catheters and tubing, polyvinylchloride tubing and films [62] , and SMP polyurethane foams [63] . After plasma sterilization (Sterrad-100St, H2O2, 52 min) PU foams are described to have an increased open porosity, but no effects on shape-memory recovery were observed [63] . Additionally, plasma treatment had no significant cytotoxic effect on L929 cells in vitro. Cytotoxicity was studied using 2- and 7-day extracts of the materials in direct contact with the cells. After 48 h, cell viability was investigated by 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl-tetrazolium bromide (MTT) colorimetric assay. Cell morphology was evaluated also. However, H2O2 residuals left on/in the materials can have a negative effect on cells. V79 fibroblast showed a strong decrease of vitality (crystal violet vital staining) after exposure to extracts of poly(methyl methacrylate) and poly(2-hydroxyethyl methacrylate) for 3 days after vapor-phase H2O2 sterilization (3 mg/l H2O2, Expert Rev. Med. Devices 7(3), (2010)



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39 min, 55°C) because of left H2O2 residuals [64] . In addition, two different acrylateL929 L929 based SMP networks (tert-butyl acrylateco-poly(ethylene-glycol)n dimethacrylate and methyl metacrylate-co-poly(ethyleneglycol)n dimethacrylate) were shown to be cytotoxic after LTP sterilization. The cell layers were still destroyed 24 h after exposure of sample extracts to the cells. The cytotoxic effect was not observed after using steam-, gas- (EO, Noxilizer), and L929 L929 L929 irradiation (g-, e-beam)-based sterilization techniques for those materials [65] . The gold standard of the LTS methods is the EO sterilization [66] . EO inactivates all organisms through alkylation of adenine and guanine residues. EO has advantages, such as adequate effectiveness at low temperatures, high penetration and compatibility with a wide range of materials. The parameters of EO sterilization Figure 2. Morphology of the L929 cells after incubation with polymer samples are time, temperature, EO concentration (A–C), as well as glass discs as negative controls (D) and copper discs as positive and moisture. For instance, polyurethanecontrols (E) in the agarose overlay test. (A) At the marginal zone almost all cells appeared as vital, red-colored cells showing the typical fusioform morphology of based SMPs can be sterilized using 10% fibroblasts. (B) The edge of the polymer sample is marked with arrows. A very high EO at 54°C, at 65% relative humidity percentage of the cells appeared vital and red-colored with fusioform morphology. for 130 min, followed by 12 h aeration at (C) The morphology of the cells underneath the polymer seems to be less clear. This is 132°C [67] . Although applicable to a wide because of the incomplete transparency of the polymeric material. Nevertheless, the red variety of materials, EO has many disadcoloration and fusion form morphology of the cells is detectable. (D) Negative controls: after incubation of the L929 cells with nontoxic glass discs, the typical fusioform vantages. EO is flammable, explosive and morphology of vital (red-colored) fibroblasts was detectable in all different zones of the can react with functional groups. Because of petri dishes. (E) Positive controls: the appearance of the cells after incubation with the inherent toxicity (mutagenicity and carcytotoxic copper discs is characterized by rounded, uncolored avital cells. The cytotoxicity cinogenicity) of EO, SMPs must be properly of the copper discs led to a change of color of the agarose layer. degassed to minimize residual EO. The influence of EO and LTP sterilization on an AB-type shape-memory polyurethane ionomers with quaternarized pyrishape-memory polymer network (N-PCLDMA-BA) composed dine moieties proved to be suited to reduce the bacterial activof PCLDMA and BA [68] was studied with L929 cells by use of ity of Klebsiella pneumoniae (-96.2%) and Staphylococcus aureus the phenol red uptake test (agarose overlay assay) (Figure 2) and the (-100%). Cytotoxicity testing by use of the agarose overlay assay MTT test, which determines the viability, activation and prolifera- with L929 mouse fibroblasts revealed no adverse effects of the tion of cells. The plasma-treated PCL/BA exhibited a significant materials on these cells [73] . toxic effect on L929 cells in both tests [69–71] . The named basic tests to investigate SMP cytotoxicity provide The biocompatibility of sterilized SMPs can be limited by general information about the influence of SMPs on cell vitality residues (e.g., endotoxins) of microorganisms that contaminated and growth behavior, and are mainly performed with immortalthe SMPs prior to sterilization. A supporting strategy to limit ized cells (e.g., V79, L929 and 3T3). Tests with organotypic and microorganism load on SMPs might be to integrate an anti regiospecific primary cells are suited to provide more detailed bacterial effect in/on the SMPs. On a shape-memory orthotic systematic information about the SMP–cell interaction in the device using shape-memory polyurethane, the anti-bacterial later medical approach. In case of SMP use as implants, the activity was achieved by blending polymer matrix with anti- process of biomaterial integration into tissues, a highly dynamic bacterial agents such as chitosan, nano-zirconia, nanozinc and interactive progress, involving numerous soluble mediators, oxide or nano-silver series. In addition, on polyurethane-based cells and components of the ECM, is a critical step. Processes SMPs synthesized using methylene diphenylene diisocyanate required for successful biomaterial tissue integration are cell (MDI), PTMO and N,N-bis(2-hydroxyethyl)isonicotinamide, migration, angiogenesis, matrix degradation and remodeling of effective nonleaching biocidal properties could be shown [72] . granulation tissue. Thereby, the tightly controlled degradation It was found that the polyurethane ionomers possessed good of the extracellular matrix is a key process that is mainly depenbactericidal activity against Staphylococcus aureus and bacterio- dent on the activity of matrix metalloproteinases (MMPs) and static activity against Escherichia coli. In addition, PCL-based the tissue inhibitors of MMP (TIMP). MMPs are capable of www.expert-reviews.com

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cleaving almost all components of the ECM, and the MMP activity is considered to be a rate-limiting step in ECM degradation. Because of this, MMP and TIMP have a critical function during biomaterial integration into hard and soft tissues. For a thermoplastic SMP PPDO-PCL composed of PPDO 38 wt% and PCL 56 wt% (PDC) [74–77] , the surface roughness was shown to have an significant influence on cell adhesion, proliferation and differentiation, as well as on the kinetics of appearance and activity levels of MMP and TIMP on primary rat cells of the upper aerodigestive tract (oral, pharyngeal and esophageal origin). The rough modification of the SMP surface supported the cell growth dependent on the cell type and the time of cell growth, rather than the smooth surface modification. 24 h after cell seeding, the different surface characteristics also resulted in an opposite pattern of the kinetics of appearance and activity levels of MMPs and TIMPs of primary epithelial cells of the upper aerodigestive tract. The highest MMP-1 and MMP-2 levels corresponded with the highest TIMP levels, and vice versa. These results were supposed to be affected by different mechanical loading by the surface roughness, and besides the chemical composition of the material surface, also by the resulting surface properties, often expressed in terms of hydrophobicity, that are the main predictors of protein adsorption and, thus, cell adhesion [77–79] . The same SMP was also tested (after EO sterilization) in direct contact with endothelial cells to provide hints on the angiogenic effects of the material [80] . The endothelial cells proved to adhere, to grow and to from confluent monolayers on the material during increasing time periods. The degree of cellular and protein adhesion, the ability to stimulate neutrophils and inflammatory cytokine production, as well as the effects on the hemostatic system, are described for two polyurethane-based SMPs, PESU and thermoset PEU [81] . Variations in PEU blood compatibility have been observed, with improved thromboresistance correlating to more alkylated soft segments [82] , higher soft segment content on the surface [83] and surface treatment, such as grafted sulfonate groups [84,85] . The PESU showed higher TNF-a production and a higher potential for macrophage and endothelial activation, respectively, compared with PEU. Incubation of both materials with human plasma for up to 4 h did not lead to any evidence of contact activation, thrombin or plasmin generation or platelet activation. In response to shear stress, no significant difference in platelet adhesion was noted. This suggests that under shear conditions, such as those seen with vascular flow, the SMPs do not enhance platelet adhesion and activation. Neuss et al. studied the effect of N-PCLDMA networks on the cell behavior of different cell types from three different species (human mesenchymal stem cells, human mesothelial cells and rat mesothelial cells) and analyzed the differentiation capacity of mesenchymal stem cells on this polymer [86] . The polymer proved biocompatibility for all tested cell types, supporting cell viability and proliferation. N-PCLDMA also supported the osteogenic and adipogenic differentiation of human mesenchymal stem cells 3 weeks after osteogenic and adipogenic induction, and the SME activation did not affect the majority of adherent cells. 364

In vivo tests for histo-compatibility & biofunctionality of shape-memory polymers

Although cells in culture represent the elementary living systems, as a very simplified system, they cannot represent the complexity of the entire organism. Hence, an in vitro system cannot replicate exactly the biodynamics of the whole human body due to the lack of possible mitigating systems (e.g., hormones, nervous system and immunity), and the lack of biotransformation and excretion pathways for their elimination in vitro. Therefore, in vivo testings in animal models are necessary prior to human clinical testing. In vivo tests are used to evaluate the local and systemic interactions (e.g., genotoxicity, carcinogenicity, reproductive toxicity, delayed-type hypersensitivity and systemic toxicity) of host tissues with implant materials or devices and their leachable/soluble/degradation components. These include physiological effects of the implant and its released components on local and systemic tissues, tissue response to the implant materials and designs, changes in extracellular matrices and regulator biomolecules and changes in the material and device due to host physiology. No single animal species represents an ideal general model for the human species. A variety of sources are devoted to the selection of appropriate animal models [87–89] . Anatomical, biochemical, physiological, pathological and/or psychological characteristics must be considered when choosing an animal model. Guidelines to carry out animal biocompatibility tests are described by standards organizations of regulatory agencies such as ASTM, ISO, NIH, FDA and US and European Pharmacopeia. Generally, in vivo tests can be divided into functional and nonfunctional tests. Nonfunctional tests are usually conducted first to study the direct interactions of the implant with the physio logical environment. In these tests, it is assumed that acute toxic or inflammatory reactions are nonspecific, and thus may be evaluated in soft tissues (e.g., subcutaneous, intramuscular and intraperitoneal), because implantation in soft tissue requires minor surgery. After evaluation in the subcutaneous tissue, the next step is the selection of an animal model and implant site for the SMP device or prototype to simulate the site that ultimately will be used in humans (i.e., functional testing). The incorporation of a biomaterial in vivo is generally assumed to require regular wound-healing mechanisms. The process of wound healing is a highly dynamic and interactive progress, which involves numerous soluble mediators, cells and components of the ECM [90] . A decisive step during regular wound healing is the angiogenesis, which is regulated by numerous growth factors, proteolytic enzymes, ECM components, cell adhesion molecules and vasoactive factors [91] . The interaction between a biomaterial and surrounding tissue requires sufficient capillary and vessel formation to ensure the transport between the tissue and implanted biomaterials [92] , and a sufficient clearance of the biomaterial–tissue–interface, for example, from degradation products. Until now, inadequate transport between the microvasculature and implanted materials has limited the utility of a variety of biomedical applications, including tissue engineering [93] . Wear particle accumulation at the biomaterial–tissue–interface was described as ‘particle disease’, and caused concerns regarding the introduction of degradable polymers Expert Rev. Med. Devices 7(3), (2010)


C

D


µm 30 µm 330 0 µm

30 µm µm 30

Figure 3. Microcirculation close to a polymer–tissue interface, which developed within 3 weeks after implantation. Newly formed microvessels 10 µm above a PDC implant [20] (A) 3 weeks after subcutaneous implantation in the rat neck; non-soluble degradation products yellow stained; micro-blood vessels visualized by fibronectin staining in green; (B) crossection, PDC red [97] .

in medical therapies [94–96] . In in vivo and in vitro experiments, the degradation behavior of the biodegradable multiblock copolymer SMP PDC, containing PPDO hard segments and crystallizable PCL switching segments, was shown [20] , synthesized via co-condensation of two oligomeric macrodiols with an aliphatic diisocyanate as a junction unit [97,98] . The study demonstrated that the in vivo degradation process of PDC is characterized by a release of microparticles, which were supposed to primarily consist of PPDO, and which didn’t negatively affect a strong integration of the SMP in the subcutis of rats [97] and mice [99] . This might be explained by the angiogenic effect of the PDC, which was shown in the hen’s egg chorioallantoic membrane test (HET-CAM-test) [71] , and also after subcutaneous implantation in rats [97] (Figure 3) and mice [99] . A SMP, N-PCLDMA-BA, based on PCLDMA and BA, was able to induce angiogenesis and strong tissue integration in male Naval Medical Research Institute mice 1 week after subcutaneous implantation [47] . The same SMP proved its capability for autoinduced regeneration of a radical stomach wall defect in rats [75,76] . No gas leakage after gas insufflation could be detected, and fast and unfavorable degradation of the polymer did not occur. A tight connection between the polymer and the adjacent stomach was found, resulting in adequate mechanical stability under the extreme pathophysical conditions of the stomach milieu (Figure 4) . Medical devices based on SMPs in development

During the past decade, the field of shape-memory materials for biomedical applications has grown into a very active area of research. This may be due to a strong clinical trend towards minimally invasive techniques [100] , which would benefit from the implantation of initially small objects that unfold to eventually bulky devices of the desired shape and functionality. This functionality often relies on mechanical forces as developed during shape recovery – for example, for wound closure by SMP suture materials or for opening of blocked vessels by stents. A summary of medical devices in development that completely or partially consist of SMPs is given in Table 1 [20,26,27,46,101–122,201–225] . As is common in a new field of research with a strong interest not www.expert-reviews.com

Review

only in academia, but also in the industry, at least some of the potential applications have so far only been briefly described in patent applications. Therefore, compared with previous reviews of the field [123,124] , this paper also summarizes information that can solely be found in the patent literature. Depending on the therapeutical aim and the required timeframe of action, biodegradable or biostable SMPs were evaluated in the respective devices. In the following, the involved concepts and set-ups of SMP-based medical devices will be discussed in sections according to their application, rather than the involved material concept or shape of the device. General surgical applications

In surgery, proper closure of wounds is required in order to allow tissue regeneration and avoid the formation of scars. While other strategies, such as surgical glue from polyalkylcyanoacrylates, might be useful in some cases, the standard procedure remains to join the edges of a wound by numerous stitches and surgical knots in a time-consuming process. Additionally, a defined pressure on the wound is advantageous for optimal healing. Finally, after healing, the removal of the suture thread by biodegradation would be desired. This aim has been addressed by the development of biodegradable SMP materials [20] . In this study, threads from thermoplastic block copolymer were shown to possess capability of shape-memory and biodegradability, as well as wound closure upon heating to moderate temperatures (Figure 5) . Another application of SMP molds has recently been suggested for the fixation of tissues or medical devices during surgery [203] . The concept involves needles of temporary straight shape that

Figure 4. Autoinduced tissue regeneration after polymer implantation in a radical rat stomach defect. (A) Radical defect (diameter: 100 mm) in a rat stomach wall 4 weeks after defect closure with a PCL/BA-based shape-memory polymers disc (diameter: 10 mm, thickness: 200 µm [47] ; the implant was fixed by conventional suturing). (B) Hematoxylin–eosin-stained histological section of the defect zone: the critical stomach wall defect was completely regenerated 4 weeks after defect closure with the shape-memory polymers [70,75] .

20°C

37°C

41°C

Figure 5. Degradable shape-memory polymer suture for wound closure that shrinks upon heating in order to join the edges of the wound for optimal healing. Reprinted with permission from AAAS, USA from [20] .

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Table 1. Potential applications of shape-memory polymer-based medical devices. SMP device

Principle

Ref.

General surgical applications Self-tightening suture

SMP thread capable of forming a self-tightening knot

[20,201]

Self-retaining suture

SMP thread that deploys spikes, anchors in tissue and holds together wounds without the need for knots

[202]

Surgical fastener

Biodegradable or biostable SMP needle; after piercing implanted non-SMP devices (e.g., hernia mesh) and subjacent tissue, temperature-induced shape recovery of the SMP needle to a curved or helical permanent shape occurs, which secures devices to the tissue.

[203]


Intravascular applications & vascular access Stents for intraluminal prothesis

Hollow devices that are at least partially radially distensible for intraluminal prothesis; different constructions and materials including biodegradable and biostable SMPs depending on the application

Aneurysm treatment

SMP devices (foams, fibers) to fill distended balloon-like dilations (aneurysms) of blood vessels or SMP prostheses that line the inner side of the vessel

Clot removal

Corkscrew-shape SMP or hybrid SMP-nitinol devices for endovascular removal of blood clots in stroke patients; devices of temporary straight shape are pierced through the clot and capture the clot upon recovery (e.g., activated by laser light or electro-resistive [Joule] heating)

[112–119]

Adjustable prosthetic valve

Tubular structure with rings of adjustable size from shape-memory alloys or SMP at both ends; the prosthetic valve implants, e.g., an prosthetic aortic valve, is secured in the aorta by ring expansion to give an hour-glass-shaped tubular structure

[212]

Catheter with microgripper

Microgripper with jaws of laser-activated SMP that can release devices intravasularly after positioning by catheter

Catheter for retrieval of devices

Catheter tip and intravascularly deposited microdevices with mating SMP patches for retrieval or repositioning

[214]

Dialysis needle

Dialysis needle with a SMP tubing for intravascular deposition; the SMP tubing in its deployed shape (larger distal diameter) reduces hemodynamic stress in arteriovenous grafts during dialysis and might prevent graft failures

[120]

Catheter for cell collection by apharesis

Temperature-induced unfolding SMP lamellae at the end of the catheter end form a funnel-like structure for improved cell collection and reduced risk of occlusion by contact of opening with the vessel wall

[215]

Anchoring cannula

Ring of biodegradable or biostable SMP around a cannula; temperature-induced expansion in an opening of an incision of the body anchors cannula

[216]

[26,103–108,204–207]

[109–113,208–211]

[119,213]

Urogenital applications [217]

Retrieval devices

Light-activated SMP devices of different permanent shapes (baskets, corkscrews) to capture concretions in the cavities of kidney, pancreas or gallbladder

Fertility control

SMP occlusion plug inserted via catheter transcervically into fallopian tube and expanded upon application of heat or light for permanent or temporary sterilization

[206,218,219]

Stents for intraluminal prothesis

Radially distensible hollow SMP devices for intraluminal prothesis; in, for example, the urinary tract; different constructions and materials including biodegradable and biostable SMPs depending on the application

[46,206,207,220,221]

Ophthalmic applications Implants for ocular hypertension treatment

SMP implants of several shapes for reduction of intraocular pressure; outflow is improved by implants that either provide interior flow pathways (stents) or modify the properties of soft tissue around physiological outflow pathways

[222,223]

Applications in the brain Neuronal electrode

SMP microactuator with enclosed conductor; very slow shape recovery for slow insertion into brain tissue with reduced astrocytic scarring; softened SMP after insertion may advantageously match micromechanics of brain tissue

[121]

SCL: Super-cooled liquids; SMP: Shape-memory polymer.

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Table 1. Potential applications of shape-memory polymer-based medical devices (cont.). SMP device

Principle

Ref.

Orthopedics & orthodontics Fixation of damaged ligament or tendon

SMP rod for screw replacement in surgical fixation of damaged ligament or tendon to bones; rod expands upon shape recovery in a bone tunnel and fixes the loose end of a ligament by defined pressure against the tunnel wall

Archwires in orthodontics

SMP wire for correction of misaligned teeth with multibracket braces; shape recovery creates forces that push the tooth in the desired position

[27]

[122]


Other internal applications [206,207,220,221]

Stents for intraluminal prothesis

Radially distensible hollow SMP devices for intraluminal prothesis; in, e.g., esophagus, bronchi, colon, bilary tract; different constructions and materials including biodegradable and biostable SMPs depending on the application

Stomach fillers for obesity treatment

SMP device that unfolds in the stomach and thus reduces the volume of the stomach for weight reduction of overweight patients

[224]

SMP device filled with SCL heat from mechanically triggered crystallization of SCL induces shape recovery and adaptation of heat package to the respective contours

[225]

External applications Heat pack

SCL: Super-cooled liquids; SMP: Shape-memory polymer.

are pierced through the implanted device, such as a hernia mesh, and the respective subjacent tissue. Upon heating to body temperature, both ends of the needle recover to a helical shape, thus fixing the implant in the desired place. Vascular applications

Major targets for the evaluation of both biodegradable and nondegradable SMPs in medical devices are cardiovascular applications. Among these, stents [101] might be the most prominent example. Based on an acrylate chemistry with PEGDMA crosslinkers, a family of AB networks has been systematically studied as a stent matrix in a report on temperature-induced SMPs [26] . Incorporation of higher amounts of the cross-linker into the network, or reducing the Tg of the switching segment, resulted in faster shape recovery of stent prototypes at body temperature in vitro. In order to allow administration by catheters, the polymer stent has been programmed to a small temporary shape by furling and rolling. Its shape recovery is shown in Figure 6. In another study, solid and perforated stents were compared, and a faster onset of recovery was found for the perforated devices [104] . In addition, tubular SMP devices were laser etched for more flexible stents that also could navigate through small tortuous vessels – however, it has been revealed that resistance to withstand the critical collapse pressure in the vessel at elevated body temperature may not be given in all cases [106] . By using triple-shape thermosensitive SMP materials, temporary stents could be designed that are implanted in their first temporary shape, expand upon heating to their required functional shape (second temporary shape) and, upon further heating, shrink for easy removal if required [102] . Besides induction of shape recovery at body temperature, indirect heating by light absorption has been evaluated as stimulus for shape recovery of temperature-sensitive polyurethane SMP stents [105] . In order to do so, the stent has been loaded with a laser-absorbing dye, crimped over a light diffuser at the end www.expert-reviews.com

of an optical fiber and exposed to near-infrared light in a synthetic artery model set-up with circulating water at 37°C. While working in principle, when setting the water flow rates to values relevant for in vivo applications, the required temperature and, therefore, a complete shape recovery, has not yet been obtained, even at laser powers up to approximately 8 W due to cooling by the surrounding medium. In addition to some other suggested applications (see Table 1), further important fields of presently developed SMP applications are aneurysm treatment and clot removal devices. Aneurysms are balloon-like dilations of vessels due to alterations of the vessel wall with highest prevalence in certain regions of brain arteries and the aorta. They may be associated with potentially lethal rupture or uncontrolled aggregation of blood clots, which could embolize peripheral tissue when removed from the aneurysm into the bloodstream. Among different treatment options, commonly used or suggested strategies that rely on endovascular implantation of medical devices include the deposition of a metallic coil (e.g., Pt) in intracranial aneurysms for a controlled thrombus formation, and subsequent closure of the aneurysm [125,126] or the implantation of vessel prostheses (stents) in aortic aneurysm [127,128] . By using polymers instead of Pt coils, an improved cell adhesion may be achievable [108], since polymers can be tailored to preferentially allow adhesion of certain cell types by the selection of building blocks for the polymer. Thus, the formation of a neointima, particularly an intact vessel wall with a functional layer of endothelial cells, may be realizable in the future and could lead to a reduction of disadvantageous recanalization of treated aneurysms. Therefore, SMP strings for polymeric coils [108,211] and SMP foams [107,109,110,208,210] have been evaluated in different set-ups, including computational simulation on vascular dynamics [109], as well as pilot in vivo studies [107] . From these first data, SMP foams can be considered to be a promising treatment option of intracranial aneurysm, since they may reduce the risk of intraoperative rupture 367


Review


[110] . Preliminary in vivo data in dogs for aneurysms of the common Urogenital applications carotid artery showed a successful closure, at least macroscopically, Besides stents such as ureteral stents [46] , as described in the in this well-healing vessel model, although no information on the section entitled ‘Regulatory aspects and potential applications cell types of the neointima was provided (Figure 7) [107] . For aortic of drug-loaded SMPs’ in larger detail, a hypothetical application aneurysms, a SMP prosthesis has been suggested [209] . of SMPs in the urinary tract uses the shape-memory concept to For the treatment of cerebral ischemia in stroke patients, mechan- capture concretions like renal calculi. During lithotripsy treatical removal of intravascular blood clots has been advantageous at ment, such stones are destroyed into small fragments that are timepoints several hours after the onset of the stroke, where clot- supposed to be passed out of the body with the urine. However, dissolving standard therapies with, for example, tissue plasminogen in some cases larger fragments may painfully block urine flow activator, are no longer effective. Accordingly, FDA approval was or even migrate back into the kidney. Such fragments can be received in the USA for a microcatheter system that is guided to the captured during urethroscopy with basket-shaped devices. It occluded vessel and passed beyond the thrombus. A nitinol wire is has been suggested that baskets or corkscrews from SMPs are advanced through the catheter, deploys at the end of the catheter in employed for this purpose [217] . a spring-like manner to a helical corkscrew shape, and captures the In another suggested application for contraception in women, clot [129] . Similar to this strategy, a corkscrew system [112] , as well SMP implants are inserted temporarily or permanently into the as an umbrella device [113] , have been designed using commercially fallopian tube by a catheter in a transcervical procedure [206,218,219] . available covalent thermosensitive SMP networks [130] . At body temperature, the implant deploys to, for example, a bulbIn further detailed studies with thermoplastic SMPs, indirect like shape, and acts as a plug that inhibits ovum migration from the heating by laser with an optimized coupling of laser light into a overy to the uterus. The efficiency, compatibility and acceptance dye-loaded SMP corkscrew has in principle been proven as suitable of such strategies remain to be shown. to induce shape recovery in air [114] . However, as is the case for all systems that employ activation by lasers or other sources of external Miscellaneous internal applications heating, a substantial cooling of the device by circulating blood In addition to medical applications with several described examin vivo requires high laser energies, at least for common commer- ples of potential SMP medical devices as discussed above, other cial SMPs with Tg’s in the range of 55–65°C. However, at the same reports cover applications in several different parts of the body and time, energy input also causes heating and potential damages to rely either on simple or complicated device architectures. Hollow blood cells and the surrounding tissue. In simulations at medium SMP tubing as described before might be used as stent also in, laser powers of 4.5 W (i.e., approximately half of what has been for example, the esophagus [206,207,220,221] . Complicated branched barely enough in an aforementioned study for stent deployment implants with internal flow pathways (stents) or spiky rods have under physiological flow conditions [105]), the energy input already at least been suggested to reduce intraocular pressure by improved resulted in a temperature increase of the surrounding blood to 49°C, which in this case was concluded to be in a safe range [115] . t=0 t=2s t = 10 s Thus, it might be useful to customize polymers to a lower switching temperature in a narrow range of 45–50°C, rather than using materials of relatively high Tg in order to enable activation at lower laser power [116] . Recently, a hybrid system consisting of a permanently corkscrew-shaped nitinol wire, thin insulated copper wires connected to each end of the nitinol (one of the copper wires was wound back around the nitinol so that both of them emerge from the same end of the nitinol) and a final SMP coating has been developed [117] . By using the SMP properties, the device was fixed in a straight shape. Due to Joule heating by t = 60 s t = 20 s t = 100 s applying current to the copper wires, the SMP switching temperature is exceeded and superelastic recovery of the nitinol Figure 6. Insertion of a transparent shape-memory polymer stent through a is enabled. After turning off the current, 22 mm catheter (18 Fr.) into a 37°C water bath induces recovery to the the SMP vitrifies, enforces the nitinol and permanent unfolded shape (black rings drawn on outer surface for better visualization). allows clot removal in vivo without the need Reprinted from [26] with permission from Elsevier. © (2007). of otherwise common guide wires [118] .

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outflow of intraocular fluid [222,223] , as relevant in certain cohorts of glaucoma patients. Unfolding of SMP devices in the stomach was suggested for application in overweight patients, where a reduction of the stomach’s volume results in a faster feeling of repletion during ingestion [224] . Shape recovery produces forces that are typically required to enable the biomedical functionality of SMP devices. Often, a fast shape recovery may be useful when immediate fixation of a device or tissue is required during an intraoperative scenario. Figure 7. Closure of common carotid aneurysms after implantation of Besides several examples discussed above, shape-memory polymers foams. (A) Macroscopic view of aneurysm neck from the this is also true for a suggested application intraluminal side of the artery. (B) Macroscopic view of a cross-section of a treated aneurysm. to fix ruptured tendons or ligaments to Reprinted from [107] with permission from Elsevier. © 2003. bones. In this application, shape recovery from a thin stretched rod to a more compact-shaped rod allows fixation of ligaments in bone tunnels, established in some cases in these first studies, depending on the in which the rod’s radial recovery forces depend on the material drug physicochemical properties and the network architecture. composition, such as the cross-linker content [27] . Moreover, This has led to multifunctional materials for biomedical applicaby employing SMP devices rather than common metal screws, tions that combine biodegradability, controlled drug release and damage of ligaments by the screw threads may be avoided dur- shape-memory capability. ing fixation. Adding a new capability to a device is not trivial, since it is not In contrast to immediate shape recovery, which is useful in unlikely that this may impair an already established functionality. several applications, slow shape recovery has been realized with a For example, incorporation of drugs may result in a disturbance neuronal probe microactuator from an amorphous SMP [121] . This of the polymer morphology or thermomechanical properties, and slow movement is required to avoid insertion-induced tissue dam- thus in an impaired shape-memory of a temperature-induced SMP. age that, among other things, leads to the formation of astrocyte Therefore, it is of high relevance to establish independence of the scars as insulating layers around the electrodes. Programming of SMP functionalities [134] . After the aforementioned general proof the 1.5-mm long device has been conducted at 50°C, which is of principle of drug-releasing SMPs, a detailed analysis of the interclearly below the polymers Tg (70°C), but also clearly above body play between drug loading, shape-memory and biodegradation will temperature, resulting in very slow movement of the probe tip at have to be conducted in the future in order to obtain tailored prop37°C (Figure 8) . Additionally, the mechanical incompatibility of erties of a device for a specific application. As will be discussed in brain tissue and stiff neuronal probes may be overcome by soften- the following, SMPs are a new promising strategy for an intelligent ing of the SMP after completed insertion. This may happen as matrix material for drug-release applications. a consequence of both warming up close to or above the SMP’s Tg, and the plasticization of SMPs by water in the physiological The concept of controlled drug release from biodegradable polymers environment [131] . Another suggested application, where persistence of recovery The concept of a controlled and sustained release of drugs from forces is required over a long time and shape recovery should biodegradable implants was developed more than 30 years be slow, is an orthodontic application of SMP threads as arch- ago [134,135] . The term ‘controlled release’ is used for formulations wires [122] . In the present example with a thermoplastic SMP that are able to deliver incorporated drugs in a controlled manner, containing PCL building blocks, a maintenance of the wire’s preferentially at predefined rates. It is aimed to tune the rates of recovery forces over months was observed at least in a water-free drug release during formulation development in order to fit the environment. However, in a dental arch model, so far only a much requirements of a specific therapeutic application. Depending on too fast shape recovery has been realized that would not allow the application, the release can be adjusted to be, for example, proper remodeling of bone around the roots of the teeth. fast, slow or pulsatile. In contrast to daily peroral medication, for example, with tablets, implants can reduce the frequency of Potential of SMP as intelligent matrix material for administration, and provide constant levels of bioactive molecules controlled drug-release systems either locally or systemically over an extended period of time (susIn order to fully explore potential applications of SMPs in the tained release). Besides ensuring delivery of the correct amount biomedical field, controlled drug-release capability of SMP of drug, a controlled-release formulation may prevent drug degmatrices has recently been studied [46,131,132] . As reviewed in radation, reduce side effects, or improve patients’ adherence to detail [133] , drug release rates over weeks or months could be the treatment [136] . www.expert-reviews.com

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Tissue

Damage

Substrate


As-fabricated

Stored

Deployed

Displacement (mm)

500 400 300 200 100 0

0

10

20

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40

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Figure 8. Slowly deploying shape-memory neuronal probe from epoxy-based shape-memory polymers (Tg : 70°C). (A) Mode of action of shape-memory polymers microactuator with enclosed conductor (central red line). (B) Light microscopy images of demonstrator in its different shapes. (C) Movement of probe tip at 37°C after programming at 50°C. Taken with permission from [121] . © 2006, Institute of Physics Publishing.

In order to tailor drug-release rates, at least the major involved processes should be well understood. Generally, dissolved drug molecules can diffuse with a certain rate out of a polymer matrix. This rate depends on several factors, including the physico chemical properties of the drug, such as molecular weight or solubility, the polymer properties, such as hydrophobicity, morphology (e.g., amorphous) and physical state (e.g., rubbery or glassy), swellability in the release medium, the polymer-free 370

volume and, if applicable, the cross-linking density, potential drug–polymer interactions and device properties such as matrix size and porosity. For drugs with relevant diffusion coefficients, a solely diffusioncontrolled drug release can be realized. It should be pointed out that the diffusion of water from the release medium into the polymer matrix is essential to initiate drug release. Typically this is the case when the dry implant matrix is transferred into a physiological environment. Following this, small-molecule drugs, at least those with some extent of water solubility, get dissolved and typically can diffuse through the matrix in order to be released. In the preferred case of SMPs with the capability to undergo biodegradation, substantial alterations of polymer properties will occur after exposure to a physiological environment. As a consequence of polymer degradation, implants lose structural integrity – that is, they undergo erosion. In principle, this erosion can either occur at the surface of the material (surface-eroding) or in the entire matrix (bulk-eroding). For example, for hydrolytic cleavage as the most commonly followed biodegradation strategy, the erosion mechanism depends on the combination of both the rate of water uptake into the matrix and the rate of hydrolytic polymer cleavage. Depending on the speed and pathway of polymer erosion, drug release and degradation/erosion may overlap. In this common scenario, diffusion and erosion both contribute to the drug’s release profile. A solely erosion-controlled drug release can be observed for molecules that do not possess diffusibility in the polymer matrix; for example, for sterical reasons due to a high molecular weight. Relevance of SMP architecture & morphology for drug-release applications

During the evaluation of drug-loaded SMPs, two steps of major importance are to load drugs into the SMP and to study and tailor the release of bioactive molecules. In principle, there are several different methods to incorporate drugs in a polymeric matrix; however, their applicability will depend on the polymer properties. Common strategies include casting of drug/polymer solutions, processing polymers in the molten state with dissolved or suspended drug, swelling of preformed networks or cross-linking a macromolecular network precursor in the presence of dispersed drug. Expert Rev. Med. Devices 7(3), (2010)


Cumulative release (%)

In this context it is important to discuss 100 potential effects of SMP properties on drug loading and release. The chemical compo80 sition of the polymer, the nature of bonds 60 in the polymer and their sensitivity to bio O H HC degradation, such as hydrolytic cleavage, N 40 N may affect the timeframe and pathway of N matrix erosion and potentially induce eroN O 20 sion-controlled drug release as discussed C CH 0 above. In addition, noncovalent interac0 5 10 15 20 25 30 35 tions, such as ionic bonds or hydrophobic 100 µm Time (day) interactions between drug and polymer, may take place when dissolved drug molFigure 9. Characterization of a theophylline-loaded shape-memory polymers ecules are brought into contact with the network derived from branched oCl octole with an average degree of polymer during drug loading. Such interpolymerization (DP) ~10 (●) or DP ~20 (○). (A) Cross-section of ● with 10 wt.% drug loading as observed by SEM. (B) Release of theophylline in PBS at pH 7.4 and 37°C. actions may increase drug solubility in Inset: Structure of theophylline. the polymer. This, on the one hand, may Adapted with permission from [132] . © 2009 American Chemical Society. increase maximum loading, for example, by less flexible in terms of modifying degradation rates by varying swelling techniques, but, on the other hand, may result in incomplete drug release due to trapping of drug in co-monomer ratios and crystallinity, since the crystalline phase needs to be preserved in order to maintain SMP functionality. the polymer phase. In order to exhibit a SME, polymer networks need to be established that can be obtained either by physical netpoints or cova- SMPs investigated as drug-loaded matrix material lent cross-linking of the polymer chains. Thermoplastic materials So far, controlled drug release has been established for covalent SMP with physical cross-links bear the big advantage of being suitable networks that were based on different chemistries: PCGDMA as for processing into a device, implant or other drug carrier of a spe- telechelics that were cross-linked by photopolymerization [46,137] , cific shape by numerous techniques in the dissolved or, if stable, oligo[(rac-lactide)-co-glycolide]tetroles (oLG) cross-linked by lowin the molten state [24,226–229] . Drug loading will typically be molecular-weight aliphatic diisocyanates [131,138] , and branched time-independent from polymer synthesis and could be carried oligo(e-caprolactone)octols (oCl) cross-linked by low-molecularout with different loading levels. weight aliphatic diisocyanates [132] . These networks were of amorDrug loading of covalently cross-linked SMP networks can be phous (oLG derived) and semi-crystalline (PCGDMA and oCl conducted by swelling the network in drug solutions with sub- derived) morphology. sequent drying. This technique has limitations in the maximum The absence of any impact of drug incorporation on the therpossible drug loading, depending on the polymer swellability, mal properties of the semicrystalline polymer network has clearly drug solubility and drug–polymer interactions. Cross-linking been shown for a library of PCGDMA-derived networks [137] . macromolecular network precursor in the presence of dispersed However, drug-loading levels of this network as obtained by the drug might possibly cause chemical alterations of drug molecules swelling technique remained low (

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