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REVIEW

1700694 (1 of 16) © 2017 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

Biodegradable Shape Memory Polymers in Medicine

Gregory I. Peterson, Andrey V. Dobrynin, and Matthew L. Becker*

DOI: 10.1002/adhm.201700694

1. Introduction

Shape memory materials (SMMs) have emerged as an impor-tant class of stimuli-responsive materials. The defining char-acteristic of SMMs is that they can be programmed into a temporary shape, and upon application of a stimulus, return to a permanent shape. The diversity of material properties and stimuli capable of triggering the shape change have made SMMs ideal candidates for various applications in the medical field. Shape memory alloys (SMAs) and shape memory poly-mers (SMPs) represent the two primary classes of SMMs.

When considering the suitability of a SMM for a medical application, there are basic design elements that must be taken into consideration, including the trigger mechanism for shape transformation, biocompatibility, method of sterilization, shape memory performance, biodegradability, and mechanical prop-erties (Figure 1).[1] In the context of implantable materials, most shape changes are designed to occur after implantation, thus shape recovery must be triggered by a stimulus from the in vivo environment or from an external source. The material and its degradation products must be nontoxic. Sterilization of the material prior to implantation is also mandatory and can be challenging with certain SMMs.[2–4] The material must possess

Shape memory materials have emerged as an important class of materials in medicine due to their ability to change shape in response to a specific stimulus, enabling the simplification of medical procedures, use of minimally invasive techniques, and access to new treatment modalities. Shape memory polymers, in particular, are well suited for such applications given their excellent shape memory performance, tunable materials properties, minimal toxicity, and potential for biodegradation and resorption. This review provides an overview of biodegradable shape memory polymers that have been used in medical applications. The majority of biodegradable shape memory polymers are based on thermally responsive polyesters or polymers that contain hydro-lyzable ester linkages. These materials have been targeted for use in applica-tions pertaining to embolization, drug delivery, stents, tissue engineering, and wound closure. The development of biodegradable shape memory polymers with unique properties or responsiveness to novel stimuli has the potential to facilitate the optimization and development of new medical applications.

Shape Memory Polymers

Dr. G. I. Peterson, Prof. A. V. Dobrynin, Prof. M. L. BeckerThe University of AkronDepartment of Polymer ScienceAkron, OH 44325-3909, USAE-mail: [email protected]

The ORCID identification number(s) for the author(s) of this article can be found under https://doi.org/10.1002/adhm.201700694.

the required shape memory performance in vivo, such as the ability to fully recover the desired shape in a specific amount of time. The permanence of the implant and the length of time a material is required to maintain its mechanical properties will determine the extent of biodegradation required. The mechanical properties of the material must be compatible with sur-rounding tissues so as not to damage the tissue or cause acute inflammation.

While the choice of material type will depend on its intended application, there are several advantages of SMPs over their SMA counterparts. SMPs are generally less dense, lightweight, easy to process, and lower cost than SMAs.[5,6] In terms of performance, SMPs can undergo larger deformation strains in the programming step (>500%) and can exhibit recoverable deformations nearly 100 times greater

than SMAs.[7] SMAs are typically stiff and the tunability of their mechanical properties is limited,[8] whereas SMP mechanical properties can be tuned over a wide range to meet the needs of a desired application. SMPs have also been shown to exhibit biocompatibility (the material and its degradation products are non-toxic) in numerous biological systems. The biocompat-ibility of some SMAs has been called into question.[9] One of the most important advantages of SMPs is their potential to be biodegradable.

Biodegradable shape memory polymers (BSMPs) have expanded the scope and clinical utility of SMPs in medical applications. In general, SMPs are primarily used for their ability to facilitate minimally invasive procedures. That is, a small object can be inserted into the body and then change its shape into a larger functional device. The primary benefit of biodegradability in these types of applications is that secondary procedures, which would be required to remove implanted materials that do not degrade, can be avoided. This is impor-tant because it helps simplify treatment and has the potential to improve procedure outcomes, decrease costs, etc. For appli-cations in which the SMP spends only a short amount of time in the body, such as an endovascular thrombectomy device (the SMP travels through a ca. 1–2 mm diameter catheter, changes its shape to mechanical retrieve the clot, then is removed from the body),[10,11] use of a BSMP would not provide significant advantage. However, the ability for a material to degrade does enable some applications (such as drug delivery, vide infra) and can be beneficial to the healing process (such as in tissue engineering).

The diversity of applications to which BSMPs can be applied is a result of the wide range of polymers that exhibit shape memory behavior. The two primary requirements for a polymer

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to possess shape memory behavior are to have two types of cross-links: 1) permanent cross-links that form a permanent net-work structure, and 2) temporary (reversible) physical cross-links that form a temporary network structure. Permanent cross-links include chemical cross-links or physical cross-links that are con-sidered to be permanent on the experimental time scale, such as microcrystalline or glassy domains (i.e., thermoplastic elasto-mers). The permanent network establishes the permanent shape of an object. Deformation of the permanent network leads to storage of an elastic energy that drives shape recovery. Tempo-rary physical cross-links include ionic bonds, hydrogen bonds, and microcrystalline or glassy domains when the operational temperature range is covering the glass transition (Tg) or melting (Tm) temperatures. The temporary network enables locking-in a new shape after sample deformation and can be reversed by changing external conditions, resulting in shape recovery. In Figure 2 we illustrate the shape programming and recovery pro-cesses by schematically showing the evolution of two networks. The original shape (Figure 2a) is deformed (Figure 2b) in com-bination with changing the external conditions (temperature, pH, salt concentration, etc., depending on the type of temporary cross-links present in the system) resulting in rearrangement of the temporary (reversible) physical cross-links (Figure 2c). Reversing the change in external conditions locks-in the distri-bution of the temporary physical cross-links and provides the temporary shape (Figure 2d). Shape recovery, triggered by reap-plication of change in the external conditions, is driven by the relaxation of stress stored in the permanent network through reconfiguration of the temporary physical cross-links (Figure 2e). After the reconfiguration of temporary cross-links is complete, the sample returns to its original shape (Figure 2a).

Temperature is one of the most common triggers for shape recovery in polymeric materials (i.e., thermal SMPs). Upon cooling, most polymers exhibit a thermal transition (Tg or Tm) resulting in the formation of temporary physical cross-links. These cross-links can be reversed by reheating the sample above the transition temperature. Thus, most polymers can be converted to thermal SMPs simply by permanently cross-linking them. Note that thermal impetus can be applied directly or indirectly through photo-,[12–14] electro-,[12,15] or magneto-thermal transduction.[16,17] SMPs can also be engineered to be responsive to other stimuli by utilizing reversible binding groups (establishing temporary physical cross-links) that are sensitive to light or chemical stimuli.[18–20] Some of these fea-tures may be inherent or may be engineered through function-alization of the polymer with the likes of hydrogen bonding or ionic functional groups.[21–23]

While an in-depth coverage of shape memory performance is beyond the scope of this review, it is important to be familiar with two main parameters that are frequently used to describe the efficacy of shape memory programming and recovery.[24] The strain fixity (Rf) and strain recovery (Rr) parameters are defined by the following equations:

εε

= × 100%temp

load

R f (1)

ε εε ε

=−−

× 100%temp rec

load int

Rr (2)

Gregory I. Peterson received his B.S. in Chemistry from Pacific Lutheran University in 2010. There he did undergraduate research with Prof. Dean A. Waldow pre-paring multiblock copolymers via anionic polymerizations. He receieved his Ph.D. in Chemistry in 2015 at the University of Washington (Seattle) under the guidance of Andrew J. Boydston. There

his graduate research focused primarily on the study of mechano-chemically responsive polymers. He is currently a postodoctoral fellow in Matthew L. Becker’s lab at the Univeristy of Akron where he works on developing new biodegradable shape memory polymers.

Andrey V. Dobrynin is the Alan N. Gent Ohio Research Scholar and Professor of Polymer Science at the University of Akron, Akron, OH. He received B.S. (1987) and Ph.D. (1991) degrees in Polymer Physics from the Moscow Institute of Physics and Technology, Moscow, Russia. Before joining the University of Akron in summer

2015, he was a faculty member at the Institute of Materials Science, University of Connecticut (2001–2015), served as a Program Director of the Condensed Matter and Materials Theory Program, Division of the Materials Research at the National Science Foundation (2013–2015). Prof. Dobrynin is a Fellow of the American Physical Society and Member of the Connecticut Academy of Science and Engineering.

Matthew L. Becker is the W. Gerald Austen Professor of Polymer Science and Polymer Engineering at The University of Akron. He received his Ph.D. in Organic Chemistry in 2003 at Washington University in St. Louis and was a NRC postdoctoral fellow and project leader in the Polymers Division of the National Institute of Standards and Technology from 2003 to

2009. At the University of Akron, Professor Becker leads a multidisciplinary team focused on developing novel polymeric materials which address unmet medical needs in regenerative medicine and drug delivery. Dr. Becker is a Fellow of the Royal Society of Chemistry and the PMSE Division of the American Chemical Society.




where εtemp is equal to the final strain of the temporary shape after programing, εload is the maximum strain applied during programming, εrec is the strain of the recovered permanent shape (after shape recovery), and εint is equal to the initial strain of the permanent shape. These parameters are obtained via cyclic thermomechanical testing, generally via tensile elon-gation. The Rf provides an indicator of how well the SMP can maintain its programmed temporary shape and the Rr provides an indicator of how well the temporary shape can recover the

permanent shape (with 100% being perfect shape fixing or recovery).

The goal of this review is to provide an overview of the BSMPs that are being used in or developed for specific medical applications. We begin with characterizing the types of BSMPs commonly used in these applications. We next discuss the specific applications of these materials. In the last part of the review, we highlight the current challenges the field faces and where we see opportunities for further research. For in-depth

summaries of past and present developments in the field of SMPs, we direct the reader to the following general reviews.[24–31] For sum-maries pertaining to the use of SMPs in biomedical applications (not specific to bio-degradable polymers) we direct the reader to the following reviews.[1,3,6,9,32–38]

2. Biodegradable Shape Memory Polymers

Biodegradable polymers generally contain hydrolytically or enzymatically sensitive bonds.[39] Hydrolytically sensitive functional groups commonly present in SMPs include esters, amides, carbamates (urethanes), car-bonates, and ureas. Enzymatically sensi-tive polymers are often naturally occurring biopolymers, such as polysaccharides, or their derivatives, such as synthetic poly(amino acids). BSMPs derived from aliphatic poly-esters, such as poly(ε-caprolactone) (PCL), poly(lactide) (PLA), and their copolymers are frequently targeted for medical applica-tions as their biodegradability and use is well established in the primary literature and in the clinic. Here, it is important to distin-guish biodegradation from bioresorption.


Figure 1. Design considerations for SMPs in medical applications. Adapted with permission.[1] Copyright 2016, American Chemical Society.

Figure 2. Schematic representation of the dual network structure of SMPs and shape memory behavior. Permanent cross-links are shown by read beads and temporary physical cross-links are shown by two-color ellipses. a) Initial shape; b) shape programming through dual network deformation; c) rearrangement of temporary physical cross-links in the strained network in response to a change in external conditions; d) fixation of the programmed shape by the tem-porary physical cross-link network structure and by reversing the change in external conditions; e) relaxation of the temporary physical network by reapplying the change in external conditions.



PCL, for instance, undergoes macromolecular degradation on a much shorter time scale than it undergoes clearance from the body via excretion or metabolization (bioresorption).[40] Table 1 describes the composition of BSMPs used in medical applica-tions, including identification of the permanent cross-links, temporary cross-links, and triggering stimulus. The chemical structures of the polymers and their components are shown in Figure 3.

2.1. Poly(ε-caprolactone)

PCL is semi-crystalline, with a Tg of ca. −60 °C and Tm between 59 and 64 °C.[40] As the Tg is below ambient temperatures, the Tm is targeted as the thermal transition for establishing tem-porary physical cross-links. PCL without covalent cross-links was shown to have Rf values below 76% and Rr values below 89% (although the Rr could be improved by decreasing the size of the testing specimen).[41] In addition to the relatively low values of fixity and recovery parameters, the PCL was limited to programming at low temperatures. The reasoning for this behavior is that the crystallites that would be able to form the permanent cross-links necessary for securing the permanent shape are destroyed upon heating to the recovery temperature. To avoid this issue, the permanent network can be established via covalent cross-linking. A commonly employed method is to cross-link the material through acrylate end-groups.[42–45] This approach leads to materials with high Rf (100%) and Rr (93%) values.[42] Alternatively, cross-links can also be established with polymer blending, by adding a polymer with a much higher Tm so that the physical cross-links are maintained over the temper-ature range of the temporary shape transformation.[46,47] One of the limitations to these methods is that the materials’ Tm, used for shape transformations, still remains significantly above physiological temperatures. This requires triggering of shape recovery with an external stimuli in vivo, such as a magneto-thermal transduction.[16,17]

Modulation of the Tm of PCL homopolymers can be achieved by employing branch points and carefully controlling the molecular weight and distribution of the branch segments. This method enables decreasing the Tm to near physiological temperature (37 °C). Longer PCL segments lead to larger and more stable PCL crystallites and thus higher Tms.[48] Alter-natively, radically cross-linking PCL with a plasticizer (allyl alcohol) was shown to be a viable method to reduce the Tm of the material (to as low as 34 °C) while still maintaining good shape memory performance.[49] The addition of co-monomers is another commonly used tactic to disrupt the crystallinity of PCL segments (although in many cases branched PCL is still used). For instance, polyurethanes with hard and soft seg-ments (the soft segment being the PCL) have been developed with Tms near physiological temperatures.[50] This requires good miscibility of the two phases as increased phase separa-tion was shown to lead to higher Tms due to the formation of larger PCL crystalline domains. Similar effects can be achieved by the use of an interpenetrating polymer network (IPN).[51,52] An additional advantage to using copolymers is the ability to tune the mechanical and degradation properties of the material for a given application.

PCL has a slow and highly anisotropic degradation time in vivo (up to 3 years).[40] The biodegradability of PCL-based SMPs used in medical applications has primarily been characterized in terms of its hydrolytic degradation in in vitro studies. In general, the high degree of crystallinity and low water uptake prohibit access of water to the ester bonds, leading to slow deg-radation rates. PCL is thought to primarily undergo hydrolysis in amorphous domains, thus decreasing the crystallinity of the materials generally leads to faster degradation rates. The crys-tallinity can be modulated using the methods described above, such as adding cross-links, decreasing the molecular weight of PCL segments, and polymer blending.[53,54] Decreasing the crystallinity, however, was not shown to enhance enzymatic degradation in a PCL-based polyurethane, instead the content of hydrogen bonding hard domains was a more significant factor.[55] Decreasing the hydrophobicity of the SMP, either via use of copolymers with more hydrophilic monomers or blending with hydrophilic polymers or composite materials, has also led to faster degradation rates.[56–59]

2.2. Poly(lactide) and Lactide-Based Copolymers

The properties of PLA are highly dependent on its microstruc-ture. For instance, poly(l-lactide) (PLLA) is semi-crystalline and has a Tg of 60–65 °C and Tm of ca. 175 °C, whereas poly(d,l-lac-tide) (PDLLA) is amorphous and has a Tg of 55–60 °C.[39] The Tg is generally targeted as the thermal transition for establishing temporary physical cross-links in lactide-based SMPs. Given that PLLA is semi-crystalline, uncross-linked PLLA has shape memory behavior due to the crystalline regions providing a permanent physical network that prohibits chain slippage during programming.[60] Uncross-linked high molecular weight PDLLA has also been shown to exhibit good shape memory properties (Rf and Rr values up to 97 and 99%, respectively), however in this case the network structure is established by chain entanglements.[61]

As the Tg of PLA is significantly above physiological tem-peratures, in vivo activation of shape changes requires heating from an external source. The Tg of lactide-based SMPs has been decreased to physiological temperature through the use of copolymers, primarily copolymers of lactide with trime-thyl carbonate and/or glycolide. The latter, poly(lactide-co-gly-colide) (PLGA) is an important biodegradable polymer that has received significant use in the medical field.[39]

One of the disadvantages of PLLA is that the material can lose its mechanical strength in vivo in ca. 6 months due to hydro-lytic degradation, however mass loss takes much longer, with complete resorption taking between 2 and 5 years.[39] As with PCL, the degradation rate of PLA is dependent on the degree of crystallinity. The crystallinity of PLLA can be decreased with cross-linking, leading to faster degradation rates.[62] PDLA degrades significantly faster than PLLA, fully losing its strength in 1–2 months, as it is an amorphous polymer.[39] The in vitro degradation of PLGA[63–65] and poly(lactide-co-glycolide-co-trimethylene carbonate)[66–68] has also been studied and shown to be faster than that of PLLA, presumably due to decreased crystallinity and the introduction of more hydrolytically sensi-tive bonds in those polymers.





Table 1. Structures of Biodegradable SMPs used in Medical Applications.

Number Structurea) Component responsible for temporary physical cross-links (triggering stimulus)

Permanent cross-link type

Reference

01 PCL with cross-linked methacrylate end-groups PCL (thermal: Tm, 47–56 °C) covalent [42–45]

02 PCL three-arm stars with cross-linked methacrylate

end-groupsPCL (thermal: Tm, ca. 36 °C) covalent [48]

03 PCL radically cross-linked with allyl alcohol PCL (thermal: Tm, 34–56 °C) covalent [49]

04 branched PCL with cross-linked acrylate end-groups PCL (thermal: Tm, 33–43 °C) covalent [69,120,122]

05 PCL functionalized with POSS PCL (thermal) physical: crystallites (POSS) [56]

06 PCL blended with 13 PCL (thermal: Tm, ca. 37 °C) physical: crystallite (PU) [46]

07 PCL blended with PUb) PCL (thermal: Tm, 45–58 °C) physical: crystallite (PU) [47]

08 PCL with cross-linked acrylate end-groups,

PTMEG with cross-linked epoxy end-groupsPCL (thermal: Tm, 43–46 °C) covalent [51]

09 PCL radically cross-linked with PSA PCL (thermal) covalent [52]

10 ethyl cellulose with grafted PCL chains that are coupled with

MDIPCL (thermal: Tm, 37–56 °C) covalent [124]

11 PCL three-arm stars and POPD cross-linked with MDI PCL (thermal: Tm, 39–40 °C) covalent, physical:

crystallite (POPD)

[101]

12 PCL three-arm stars and PHBV cross-linked with MDI PCL (thermal: Tm, 39–40 °C) covalent, physical:

crystallite (PHBV)

[98]

13 PU prepared from MDI, BDO, and PCL PCL (thermal: Tm, 36–52 °C) physical: crystallite (hard

domains)

[50,72]

14 PCL and PDMS block copolymer with HDI linkages PCL (thermal: Tm, 49–50 °C) physical: hydrogen bonding [116]

15 PCL and AT with HDI linkages PCL (thermal: Tm, 21–51 °C) physical: hydrogen bonding [55]

16 OCL and ODX coupled with TMDI OCL (thermal: Tm, ca. 40 °C) physical (crystalline ODX) [123]

17 OCL graft copolymer (OG backbone) cross-linked with HDI OCL (thermal: Tm, 42–48 °C) covalent [54]

18 OCL dimethacrylate, cross-linked with BA as comonomer OCL (thermal: Tm, 25–52 °C) covalent [8,106]

19 radically cross-linked PCL and PEG copolymer PCL/PEG (thermal: Tm, 52–60 °C) covalent [58]

20 PCL and castor oil cross-linked with HDI or PHDI PCL/castor oil (thermal: Tg, −8–35 °C) covalent [108,110]

21 o(CL-co-GA) with cross-linked methacrylate end-groups o(CL-co-GA) (thermal: Tm, 25–55 °C) covalent [57,90]

22 o(CL-co-GA) diacrylate cross-linked with BA as comonomer o(CL-co-GA) (thermal: Tm, 25–50 °C) covalent [92]

23 PLLA PLLA (thermal: Tg)c) physical: crystallite [95,102]

24 radically cross-linked PLAd) PLA (thermal: Tg, ca. 48 °C) covalent [96]

25 PLLA six-arm stars and AT with HDI linkages PLLA (thermal: Tg, 56–58 °C) covalent, physical:

crystallite

[62]

26 PU from PDLLA and HDI and functionalized with POSS PDLLA (thermal: Tg, 48–49 °C) physical crystallites (POSS) [117,121]

27 p(DLLA-co-TMC) with cross-linked methacrylate end-groups p(DLLA-co-TMC) (thermal: Tg, 10–37 °C) covalent [125]

28 p(DLLA-co-TMC) cross-linked with PETA p(DLLA-co-TMC) (thermal: Tg, ca. 23 °C) covalent [107]

29 p(DLLA-co-TMC) p(DLLA-co-TMC) (thermal: Tg, 19–51 °C) physical: chain

entanglement

[114,115]

30 p(LLA-co-GA-co TMC) p(LLA-co-GA-co TMC) (thermal: Tg, 41–44 °C) physical: crystallite [66–68]

31 PLGAe) PLGA (thermal: Tg, 36–51 °C) physical: chain

entanglement

[65,73,97]

32 PLCL, PLGA blend PLGA (thermal: Tg, 50–60 °C) physical: crystallite (PLCL) [112]

33 PLGA nanoparticles coated with chitosan PLGA (thermal: Tg, 45–50 °C) physical [64]

34 OLGA four-arm stars, cross-linked with TMDI OLGA (thermal: Tg, 51–53 °C) covalent [63]

35 PGS functionalized with UPy moieties hydorgen bonded segments

(thermal: Tm, 60–80 °C)

covalent, physical:

hydrogen bonding

[91]

36 PGD PGD (thermal: Tm, 36–46 °C covalent [113]

37 PPC, PCL blend PPC (thermal: Tg, 33–37 °C) physical crystallites (PCL) [93]



2.3. Other Notable Materials

The number of non-caprolactone/lactide-based SMPs used in biomedical applications is limited. The lack of innovation is likely tied to the wide range of materials properties obtain-able with just those two classes of polymer (e.g. these mate-rials can be tuned for use in applications that require vastly different mechanical properties, such as vascular and bone tissues) and the large investment required to verify the bio-compatibility, degradability, mechanical properties, etc. of new polymeric compositions. After caprolactone/lactide-based SMPs, polyurethane-based SMPs are the most utilized, largely due to their synthetic flexibility, which enables tuning of the thermal, mechanical, and degradation properties of the material. A major limitation of this class of material is that the degradability of many formulations has not been fully explored and confirmed to occur on a clinically relevant timescale.

The majority of BSMPs are only responsive to thermal stimuli. Through the use of additives (e.g., dyes or metal nan-oparticles), BSMPs can be made responsive to light[14,69–71] or magnetism.[17] BSMPs can also be responsive to other

stimuli, such as high-intensity focused ultrasound (HIFU) without the use of additives.[64,72] These techniques, however, are energy transduction processes and still rely on heating of the material to trigger shape recovery. Similarly, solvent-based approaches can be used in which the polymer is plasticized, lowering the thermal transition of the material and enabling shape recovery.[27,65,73] BSMPs that have been used in medical applications that directly respond to alternative stimuli include a pH sensitive PU functionalized with pyridine functionali-ties,[74] a hydrogel sensitive to two different ions, providing a multi-shape memory polymer (MSMP,[31] i.e. more than one temporary shape can be programmed),[75] and a hydrogel whose temporary physical cross-links are dissolved by water to trigger shape recovery.[65,76]

3. Medical Applications

The majority of proposed medical applications of BSMPs are related to embolization, drug delivery, stents, tissue engi-neering, or wound closure. Table 2 provides a list of the medical applications.


Number Structurea) Component responsible for temporary physical cross-links (triggering stimulus)

Permanent cross-link type

Reference

38 PU prepared from HPED, TEA, and HDI PU (thermal: Tg, 50–86 °C) covalent [70,71,77,83,85,126]

39 PU prepared from HPED, TEA, and TMDI PU (thermal: Tg, ca. 75–85 °C) covalent [78,86,152]

40 PU prepared from HPED, TEA, TMDI, and HDI PU (thermal: Tg, 63–75 °C) covalent [83,84,87]

41 PU prepared from PCL, HPED, TEA with HDI, IPDI, or

TMDI linkagesPU (thermal: Tg, −19–70 °C) covalent [59]

42 PU primarily derived from TMDI and BD, cross-linked with

electron beam radiationPU (thermal: Tg, 37–80 °C) covalent [104]

43 PUU prepared from PDLLA, BDA, HDI PUU (thermal) physical [147]

43 PU prepared from MDI, BDO, PTMEG PTMEG (thermal: Tg, 66–73 °C) physical [109]

45 PDLLA-PEG-PDLLA triblock copolymer PEG (thermal: Tm, ca. 50 °C) physical: phase separated

domains

[149,153]

46 PLGA-PEG-PLGA triblock copolymer PEG (thermal: Tm, ca. 50 °C) physical: phase separated

domains

[118,149]

47 PEG with cross-linked acrylate end-groups f) PEG (water) covalent, physical:

crystallites

[65,76]

48 PEG, MDI, and BIN copolymer BIN segments (pH) physical: hydrogen bonding [74]

49 copolymer of AN, AA, and BAC AN/AA (ions) covalent, physical: ionic [75]

50 gelatin functionalized with UPy gelatin (thermal: Tg, 89 °C, and water) physical: hydrogen bonding [94]

a)Additives such as nanoparticles, dyes, drugs, etc. are not listed for simplicity, however, they may have an effect on shape memory properties. b)Elastollans TPU, 1185A. c)Shape recovery was assisted by a balloon. d)The steriochemistry was not described. e)Used in a bilayer structure with PLLA.[97] f)Used as a coating over PLGA. Abbre-viations: AA = acrylic acid, AN = acrylonitrile, AT = aniline trimer, BA = n-butyl acrylate, BAC = N,N′-bis(acryloyl)cystamine, BD = 2-butene-1,4-diol, BDA = butadiamine, BDO = 1,4-butanediol, BIN = N,N-Bis(2-hydroxylethyl)isonicotinamine, CL = ε-caprolactone, DLLA = D,L-lactide, GA = glycolide, HDI = 1,6-hexamethylene diisocyanate, HPED = N,N,N′,N′-tetrakis(hydroxypropyl)ethylenediamine, IPDI = isophorone diisocyanate, IPN = interpenetrating polymer network, LLA = L-lactide, MDI = Methylene diphenyl 4,4′-diisocyanate, OCL = oligo(ε-caprolactone), ODX = oligo(p-dioxanone), OG = oligo(glycerine), OLGA = oligo(lactide-co-glycolide), PDMS = polydimethyl-siloxane, PEG = poly(ethylene glycol), PCL = poly(ε-caprolactone), PETA = pentaerythritol triacrylate, PGD = poly(glycerol dodecanoate), PGS = poly(glycerol sebacate), PHBV = poly[(R)-3-hydroxybutyrate-co-(R)-3-hydroxyvalerate], PHDI = poly(hexamethylene diisocyanate), PLCL = poly(L-lactide-co-caprolactone), PLGA = poly(lactide-co-glycolide), PLLA = poly(L-lactide), POPD = poly(2-oxepane-1,5-dione), POSS = polyhedral oligomeric silsequioxane, PPC = poly(propylene carbonate), PSA = poly(sebacic anhydride), PTMEG = poly(tetramethylene ether glycol), PU = polyurethane, PUU = poly(urethane urea), TEA = tris(2-hydroxyethyl)amine, TMC = 1,3-trimethylene car-bonate, TMDI = isomeric mixture of 2,2,4- and 2,4,4-trimethyl-1,6-hexamethylene diisocyanate, POSS = polyhedral oligomeric silsequioxane, UPy = ureido-pyrimidinone.

Table 1. Continued.




Figure 3. Chemical structures of polymers and small molecules used in the preparation of biodegradable SMPs.



3.1. Embolization

Aneurysm rupture is often severely debilitating, if not fatal. Endo-vascular treatment methods, such as occlusion with platinum coils, have become commonplace due to their minimally-invasive nature. However, occlusion success rates with platinum coils range from 15–85%,[59] due to low endothelial cell attachment, low fill volume (of coils in the aneurysm), and coil compaction, shifting, or migration.[77] Coating the platinum coils with a bio-compatible materials, such as a polymer, has shown promise in increasing treatment outcomes. For instance, by coating a metal coil with a polyurethane-based BSMP foam, rapid stable thrombus formation was achieved in in vitro and in vivo models, while significantly increasing volume occlusion values.[78] An alterna-tive approach is to fully replace the coils with a BSMP. A poly-urethane-based thermal SMP coil was deployed in a simulated aneurysm model, demonstrating the feasibility of this method.[79]

Due to the inherent risks associated with using a coil design, significant effort has been devoted toward using BSMP foams for aneurysm embolization. In this approach, a compacted foam would be delivered to the aneurysm, and the thermally triggered shape change would lead to expansion of the foam and filling of the cavity (Figure 4A). The majority of work in this area has been in the preliminary stages of application development, such as studying foam deployment with in vitro models,[70,71] studying biocompatibility,[80] tuning mechanical properties,[81] and enhancing radiocontrast.[77] While most of the foam formulations are polyurethane-based and their degra-dation behavior is largely unexplored, efforts have been made to produce foams with tunable biodegradability.[59] Preliminary in vivo studies of BSMP foams have shown promising results, such as complete vessel wall formation in a porcine aneurysm model (Figure 4B),[82] and permanent occlusion or improve-ment of angiographic scores in a canine model.[5] Endovascular


Table 2. Medical Applications and Devices Based on BSMPs.

Applicationa) Notes Polymer Type Reference

Embolization Coils for aneurysm occlusion 39, PUa) [5,59,77–79]

Foam for aneurysm occlusion 38, 40, 41, PUb) [80,82,83]

Laser-activated deployment of foam for aneurysm occlusion 38 [70,71]

Vascular occlusion 31, 38, 39, 40, 47 [65,76,84–87]

Drug Delivery Drug release independent of shape recovery 05, 09, 17, 21, 22, 30, 34, 35, 37 [52,54,56,57,63,67,68,90–93]

Drug release dependent on shape recovery 50 [94]

Drug release induced by high-intensity focused ultrasound 13, 33 [64,72]

Drug released by pH change 48 [74]

Stents Coronary (or non-specific) stent 06, 12, 23, 24, 31, 37 [46,93,95–99]

Bile duct stent 23 [102]

Intracranial stent 42, PUc) [103,104]

Ureter stent 21 [90]

Tracheal stent 01 [45]

Retractable stent PUd) [34]

Drug-eluting stent 11, 19, PUd) [58,100,101]

Tissue Engineering 2D scaffolds 01, 15, 18, 32, 36 [42,55,106,112,113]

3D scaffolds 20, 28, 31, 39, 44, 45 [73,107–110,152,153]

Bone tissue engineering 01, 14, 25, 26, 29, 46 [43,44,62,114–118]

Dynamic surfaces 02, 03, 04, 26, 49, PUd) [48,49,69,75,119–122]

Wound Closure Sutures 07, 10, 16 [47,123,124]

Staples 30, PLAe) [34,66]

Suture-less anastomosis PLAd) [34]

Annulus fibrous closure device 27 [125]

Hemostatic device 38 [126]

Other Artificial spinal disc 46 [149]

Contraception 43 [147]

Actuator 08 [51]

Orthodontics arch wire 13 [50]

Annuloplasty ring PUf) [148]

a)Calomer, a block copolymer PU. b)Cold Hibernated Elastic Memory (CHEM)-based PUs. c)MM7520 PU from DiAPLEX Company which is a segmented PU and has a microphase separated morphology. d)Further details about the chemical structure not described. e)A commercial biodegradable PLA-based staple from Insorb. f)PU resin from MCP Iberia company with reference 3174.



delivery of a foam device in a porcine model was also dem-onstrated, however further optimization is required so that smaller catheters can be used for treatment of intracranial aneurysms.[83]

Peripheral venous disorders, such as chronic venous insuf-ficiency, are characterized by abnormal blood flow through ves-sels (often in the legs) which can cause pain, dysfunction, and potentially hemorrhage. Peripheral occlusion devices can be used to block or direct blood flow to help alleviate symptoms. BSMP foams have recently been proposed for use in peripheral vascular occlusion devices.[84–87] In these applications, reticu-lated foams (the membranes between adjacent foam pores have been removed) have been used, which has been proposed to improve the blood permeability of the foam and enhance healing. While rapid occlusion was demonstrated in an in vivo

porcine model (with endovascular delivery of the SMP occlusion device, Figure 5), it was concluded that further research was needed to determine the effect of reticulation on healing.[84] For applications in which occlu-sion is only needed for a shorter time scale (a few weeks), occlusion devices were devel-oped based on hydrogels or hydrogel coated materials.[65,76] Vascular occlusion of various arteries in a rabbit model was demonstrated, although the in vivo degradation profile and recanalization (reestablishing blood flow) rate are intended to be the topics of future studies.[65]

3.2. Drug Delivery

Controlled release formulations (CRF, drug loaded polymer matrices) generally aim to improve patient quality of life by enabling spatiotemporal control of drug release.[88,89] By tuning the location, quantity, and duration of drugs in the body, their therapeutic ben-efits can be maximized. For instance, local application of a drug (e.g. implantation of a CRF near diseased tissue) can increase the local drug concentration while minimizing systemic exposure, which is important for toxic drugs. CRFs also enable tunable drug release profiles (fast, slow, pulsatile, etc.), providing control of the quantity and dura-tion of the drug exposure. In some cases, decreasing the frequency of drug adminis-tration (through sustained release profiles) simply for the purpose of increasing patient adherence to the treatment program is desired.

In BSMP-based CRFs, the BSMP gen-erally plays one of two roles: 1) its shape recovery induces or enhances drug release, or 2) its shape recovery has no influence on drug release and is solely intended to facili-tate delivery of the CRF to the body (e.g.

minimally invasive procedures).[88] The biodegradability of the BSMP can also play multiple roles. As with other medical applications, biodegradability can eliminate secondary pro-cedures needed to remove the device at the conclusion of the treatment program. Degradation can also influence the rate of drug release. As the polymer matrix acts as a diffusion barrier to slow the release of drug, release rates can increase as that barrier degrades (meaning drug release can be both diffusion and erosion controlled).

Most BSMP-based CRFs developed thus far are designed to follow the second role described above (i.e., drug release and shape recovery are independent from each other).[52,54,56,57,63,67,68,90–93] However, the presence of drug can still influence shape memory behavior. Several studies have shown that the presence of drug decreases shape memory


Figure 4. a) From left to right, the stages of a BSMP foam being deployed into an aneurysm. b) Foam deployment and healing process of a vein pouch model for aneurysm treatment at 0, 30, and 90 days. Adapted with permission.[82]



performance (i.e., decreased Rf or Rr parameters),[56,57,63] or may lead to complete loss of shape memory behavior with cer-tain drug loadings.[54] It is also possible for drugs to influence the Tg or Tm of materials and correspondingly the tempera-ture at which shape recovery occurs. There are a few examples in which drug release is dependent on the shape change. In a head-to-head comparison of a hydrogel programmed into a temporary shape and in its permanent shape, the rate of drug release was decreased in the material in the temporary shape.[94] Recovery of the permanent shape led to increase of the release rate to the same level of the permanent shape. Drug release that is enhanced during shape recovery has also been demonstrated with BSMPs that are triggered with HIFU.[64,72] Shape recovery could be turned “on” and “off” via application of HIFU, and the drug release responded in a synchronized pul-satile manner (Figure 6). Similar behavior was observed in a

pH responsive BSMP, although in this case the enhanced drug release was due to pH dependent hydration of the material.[74]

3.3. Stents

Coronary arterial stenosis (narrowing of the arteries that supply blood to the heart) can be treated with stents during percutaneous transluminal coronary angioplasty (a cath-eter-based, minimally invasive procedure that opens arteries, Figure 7). Metal stents are prone to restenosis (re-narrowing of the artery) after 6 months,[58] however, biodegrad-able polymeric stents have been identified

as potential candidates to combat this issue. One of the first BSMP stents was the Igaki-Tamai stent (made of PLLA) which required the use of a heated balloon during deployment due to its Tg being significantly higher than physiological tempera-ture.[95] The balloon enabled heating to 70 °C, where the stent could fully expand in 0.2 seconds, compared to 20 minutes at 37 °C without a balloon. Thermal-transduction processes pro-vide an alternative to direct heating. One such example is stent deployment with magneto-thermal heating.[96] In order to sim-plify stent deployment and decrease the risk of thermal injury to vessels, many applications target self-deploying stents at physiological temperatures. In general, Tg-based stents deploy on a slower scale (minutes),[97] whereas Tm-based stents have been developed that deploy in less than a minute.[46,98] The Tm is generally a sharper transition than the Tg and thus leads to faster shape recovery. Fast deployment (less than one minute)


Figure 6. a) Illustration of the concomitant shape recovery and drug release with application of high-intensity focused ultrasound (HIFU). b) Synchro-nized shape recovery and drug release controlled by HIFU. c) Photographs depicting the shape recovery. Pictures include the time of HIFU exposure and the roman numerals indicate where those shapes fall on the experimental time on the plot in (b). Adapted with permission.[64] Copyright 2013, American Chemical Society.

Figure 5. Angiograms acquired (left) before implantation of the BSMP foam vascular occlu-sion devices (VODs) and (right) after vessel occlusion. Circles indicate location of the VODs. Adapted with permission.[84] Copyright 2014, Elsevier.



is ideal as it minimizes the risk of stent migration before full expansion to the vessel wall.

As with other implantable materials, the advantage of using BSMPs is that secondary procedures to remove the stent should not be needed. For coronary stents, the stents must maintain their scaffolding strength for at least 6 months in order to overcome the stresses associated with vessel remodeling.[95] The degradation rate of most stents has not been explored in in vivo studies. One notable exception is the Igaki-Tamai stent, which was studied for ca. 10 years in human trials.[99] The PLLA stent fully degraded within 3 years and the long-term safety of the stent was supported.

Additional functionality has also been imparted to stents. Retractable stents have been demonstrated using MSMPs.[34] Certain medical complications might require repositioning or removal of implanted stents (prior to significant polymer degra-dation and tissue growth), thus the ability for a stent to retract from the vessel wall could facilitate this process. Additionally, drug-eluting BSMP stents have been developed that help pre-vent restenosis or promote healing.[58,100,101] These stents are still in the developmental stage and have yet to progress to in vivo studies. It is important to note, that while the focus of this section has been on application of stents in coronary arteries, BSMP stents are also being developed for use in other locations in the body including the urethra,[90] trachea,[45] bile ducts,[102] and intracranial arteries.[103,104] Application to other areas requires changing the specific stent design, dimensions, and delivery method.

3.4. Tissue Engineering

Tissue engineering aims to regenerate damaged tissues by devel-oping materials that serve as biological substitutes and restore, maintain, or improve tissue function.[105] Early studies on BSMPs in tissue engineering focused on exploring the biocompatibility of 2D scaffolds and the effect of shape recovery on adherent cells.[42,106] Unaffected confluent cell monolayers, sub-confluent regions, and regions containing apoptotic cells were observed after thermal shape recovery. Apoptosis was presumed to occur due to exposure to high deformation forces during recovery of large strains (>100%). The viability of 3D scaffolds has also been explored.[107–109] Porous BSMP scaffolds prepared by three-dimensional printing (3DP) were shown to have good cytocom-patibility. Recent studies on 3D scaffolds from bio degradable[110] and non-biodegradable[111] SMPs have shown shape transforma-tion to not have detrimental effects to adherent cells.

In comparison to the development of methodologies for vas-cular,[112,113] skeletal muscle,[55] and nerve tissue,[73] considerable

effort has been devoted to the use of BSMPs in bone tissue engineering. Fibrous scaffolds prepared by electrospinning have shown promising results in enhancing osteoblast prolifer-ation, alkaline phosphatase activity, and mineralization.[114–116] Similar results were obtained with 3D scaffolds from BSMP foams.[43,44] A potential advantage to these systems is the ability to conform to irregular sized defects in bone (Figure 8). BSMPs have also been utilized as support materials for bone grafts,[117] including multifunctional materials that also deliver stem cells to augment healing.[118] Importantly, the utility of several BSMP formulations was shown to be enhanced with additives or reactive moieties, without negatively influencing their shape memory behavior. For instance, polydopamine and hydroxyapa-tite have been used as additives to enhance the bioactivity of


Figure 7. Illustartion of the radial expansion of a stent in a blood vessel with stenosis. Adapted with permission.[96] Copyright 2017, American Chemical Society.

Figure 8. a) Demonstration of the maleability of PCL foam when heated above its Tm. b) Foam mechanically fitted into a model irregular defect. c) Foam removed from the defect (after cooling) showing that the new temporary shape is retained. Adapted with permission.[43] Copyright 2014, Elsevier.



BSMP formulations. Polydopamine had negligible influence on the shape memory performance of a PCL-based BSMP while enhancing osteoblast adhesion, proliferation, and osteogenic gene expression.[43] Hydroxyapatite increased alkaline phos-phatase activity and mineralization in a PLA-based BSMP and, notably, also increased the shape memory performance.[115] An electroactive PLA-based BSMP, that was polymerized with ani-line trimer functionalities, was also shown to enhance osteo-genic differentiation of C2C12 myoblasts, without loss of shape memory performance.[62] Overall, most of these materials show promise in bone tissue engineering applications, but further study, including expansion to in vivo studies, is needed.

In addition to facilitating minimally invasive procedures, BSMPs have been used in dynamic surfaces capable of con-trolling cell orientation, growth, and differentiation. The use of patterned microgrooves has been shown to enable control of cell orientation and morphology. Aligned cells cultured on microgrooves were shown to lose their alignment when shape recovery was triggered (through either direct heating or photo-thermal transduction).[48,69,119,120] The behavior of cells losing their orientation was also demonstrated in a system in which aligned nanofibers underwent shape recovery to a random orientation.[121] Control of cell orientation and morphology is essential as it has the potential to control cellular function/dif-ferentiation. A different microgroove system was developed in which the width of the grooves decreased upon heating, leading to elongated cells and enhanced myogenesis of stem cells.[49] Switching the orientation of microgrooves can also control cel-lular behavior.[122] Cardiomyocyte sheets were shown to be able to reorganize their contractile direction in response to a 90° transition in substrate pattern orientation. Rather than using surface topology, stem cell differentiation has also been shown to be controlled in a folding 3D scaffold.[75] A flat scaffold underwent shape recovery to a 3D cube shape in which cells were oriented horizontally, vertically, or up-side down. The cells on the walls and bottom of the cube were directed toward osteo-genesis whereas the up-side down cells were directed toward adipogenesis.

3.5. Wound Closure

Simple wound closure devices including sutures,[47,123,124] clamps,[34,66] and springs (for suture-less anastomosis)[34] have been demonstrated with BSMPs. In the case of sutures, their temporary shape consisted of fibers that were stretched to 200 or 1000% strain. Applying heat to the suture triggered shape recovery to the shorter length. Sutures that were loosely applied in tissue or loosely tied would then tighten around the tissue or form a knot, respectively. So called self-tightening or -tying sutures could potentially simplify the application of sutures in procedures with limited space or visibility. Exam-ples of more complicated wound closure devices include an annulus fibrosus closure device and a hemostatic device. The former was designed to seal a tear in the annulus fibrosus of intervertebral discs.[125] The biocompatibility of the device was confirmed and its minimally invasive deployment in the disc of an ex vivo canine spine was explored. The latter was designed by combining a shape memory foam, hydrogel, and iodine to

provide a prototype device that could promote hemostasis as well as uptake a large amount of fluid and prevent bacterial infections.[126]

4. Future and Outlook

4.1. Materials

Important areas for future work in the development of BSMPs include exploring new polymer classes, developing responsive-ness to new stimuli, and thorough characterization of mate-rials properties and performance in vivo. While there are a large number of BSMP formulations (Table 1), the majority of materials are derived from polyesters or polymers that contain hydrolyzable ester linkages. There are numerous biodegradable polymers (e.g., polyphosphazenes, polyanhydrides, polyacetals, and poly(ortho esters)) that have shown promise in biomedical applications,[39,127] yet their shape memory properties remain underexplored. As the majority of demonstrated BSMPs are only responsive to thermal impetus, development of materials that are responsive to other stimuli may be beneficial to appli-cations that require temporal control of shape changes, which could be triggered by specific biological signals or an external stimulus. Regardless of the materials that are selected for fur-ther use, more in-depth study of the materials properties, shape memory performance, and in vivo biodegradation (including long-term studies to determine how long it takes for materials to lose mechanical strength, clearance rates, degradation mech-anisms, etc.) is critical for identifying the applications in which these materials will be best suited and optimizing the materials for those applications.

Another important area of future work will be in expanding currently developed BSMPs for specific medical applications. Considerable research has been devoted to the development of novel BSMP formulations that have the potential for use in medicine. For recent examples we direct the reader to the fol-lowing references.[128–143] In the Becker lab, we have developed amino acid-based poly(ester urea)s (PEUs) and explored their shape memory behavior.[144,145] These materials are ideal can-didates for BSMPs as they have tunable mechanical properties, tunable biodegradation rates, excellent blood and tissue com-patibility, and do not cause inflammation from acidosis during in vivo degradation. PEUs have excellent shape memory perfor-mance (Rf and Rr values up to 99%) and also exhibit multi-shape memory and temperature memory properties. The translation of these and other materials to medical devices represents an important next step. Improving the scope of available materials should further enable optimization of current medical applica-tions and expansion of BSMPs to new applications.

4.2. Applications

The majority of demonstrated medical applications of BSMPs are in the preliminary or early developmental stages. In vivo studies to further support the viability of most of these new treatment modalities are needed. Additionally, long term studies that allow for complete degradation of the BSMP will be




important, especially given the potential for very long degradation times for some mate-rials. For embolization applications, signifi-cant device optimization is required in order to be amendable to the use of small catheters that can access any part of the body (such as is required for the treatment of intracranial aneurysms). For drug delivery applications, development of more CRFs in which drug release is dependent on shape recovery is required. Coupled with responsiveness to specific biological ques or an external stim-ulus, better spatiotemporal control of drug release may be achieved leading to increased efficacy of thera-peutic agents. For stent applications, further developing drug eluting stents and retractable stents represent two promising areas for continued research. For tissue engineering applica-tions, more work is necessary on 3D scaffolds, particularly scaf-folds in which cells have already been seeded prior to shape transformation. The continued development of dynamic 2D and 3D surfaces and scaffolds for tissue engineering is impor-tant given the promising results this far in directing cellular growth and stem cell differentiation. For wound closure appli-cations, new creative techniques or devices using BSMPs that can help simplify common medical practices and techniques are needed.

For some medical applications, the next step in the develop-ment process is to incorporate BSMPs, including replacing non-degradable SMPs with BSMPs. For example, a non-degradable polyurethane-based SMP graft was used to treat segmental bone defects.[117] The graft maintained defect stability, inte-grated with the native bone in 12 weeks post-surgery, and pro-vided torsional stability comparable to an allograft. Switching to a BSMP is essential to healing as it enables replacement of the graft material with bone. In another example, a non-degra-dable epoxy-based material was used as the anchor for neuronal probes for chronic recording of brain activity.[146] It was shown that the SMP material properties were more compatible with brain tissue than metal probes and the shape memory actuation enabled slow deployment of the probes, leading to decreased trauma. Switching to a BSMP could be useful for facilitating the removal of the probes once brain monitoring is completed.

There are several important medical applications that cur-rently use BSMPs (outside of the major applications discussed above) that could use further development. Examples include an orthodontics arch wire for aligning teeth,[50] an implantable (fallopian tube) contraception device,[147] an annuloplasty ring for the treatment of mitral valve insufficiency,[148] an actuator,[51] and an artificial spinal disc.[149] The last two examples high-light some of the unique functions achievable with SMPs that have been underexplored for medical applications. The actu-ator utilizes a two-way SMP that can switch between two dif-ferent shapes by changing between two different temperatures, without requiring new programming steps.[31] One of the pro-posed applications of the actuator was an artificial tendon. The artificial spinal disc utilizes a BSMP that undergoes hydration induced stiffening. Water induced phase separation and crystal-lization within an amphiphilic block copolymer led to mechan-ical strengthening of the material. It was proposed that this

behavior would be advantageous for spinal disc applications (Figure 9) as softer materials would be easier to manipulate and install, but after shape recovery to fill the disc space between vertebrae, the disc would stiffen for enhanced weight-bearing performance.

4.3. Emerging Areas

Other underexplored research topics that have the potential to bring exciting functionality to medical devices or methodolo-gies include biodegradable MSMPs, 4D printing, and temporal control of shape recovery. Few ideas for applications that could benefit from MSMPs have emerged. New creative concepts or device designs are needed that can take advantage of the ability of materials to undergo multiple shape transitions. 4D printing (3D printed objects that undergo shape transformations after printing) has been utilized for a few applications,[45,96,108] but is likely being limited by the small number of BSMPs that have been explored for 3D printing. The orthogonality of 3D printing to other manufacturing techniques, ease of use, generally lower cost, and degree of customizability of printed devices provide routes to highly-personalized and patient-specific therapies that have the potential to revolutionize the medical field.[150] Con-tinued development of BSMPs that are amendable to various 3D printing techniques is critical for this vision. Temporal con-trol of shape changes has, to this point, been primarily achieved through the use of external stimuli. Autonomous activation of SMMs, with predetermined rates of activation, has the potential for use in drug delivery systems, actuators, and dynamic sur-faces for driving cell fates,[151] and merging this concept with BSMPs could lead to simplifying current treatment modali-ties which could decrease medical costs and improve patient outcomes.

5. Conclusion

A large number of BSMPs have been developed and are cur-rently being targeted for use in medical applications. The chemical and structural diversity of available materials, while limited, has enabled the use of BSMPs in a wide range of appli-cations. The majority of applications simply utilize BSMPs for their ability to enable minimally invasive procedures. How-ever, considerable testing and optimization is still required before such materials will be used commonly in the clinic.


Figure 9. Demonstration of the application of a BSMP as an artificial spinal disk. a) Collapsed disk space between model invertebra. b) Insertion of artificial disk in its temporary shape. c) Shape recovery to fill the disk space. d) Stiffening of the hydrated spinal disk. Adapted with permission.[149] Copyright 2017, American Chemical Society.




Further development of new BSMPs with unique properties and reactivity, such as multi-shape memory behavior, two-way shape memory behavior, dynamic materials properties, and autonomous shape recovery may further lead to simplifica-tion of medical procedures as well as the discovery of unique treatment modalities.

AcknowledgementsThis work was supported by the Ohio Department of Development’s Innovation Platform Program and the Biomaterials Division of the National Science Foundation (DMR-1507420). MLB acknowledges support from the W. Gerald Austen Professor in Polymer Science and Polymer Engineering endowed by the Knight Foundation.

Conflict of InterestThe authors declare no conflict of interest.

Keywordsbiodegradable polymers, biomaterials, functional materials, medicine, shape memory polymers

Received: June 2, 2017Revised: July 4, 2017

Published online:

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biodegradable shape memory polymers in medicine€¦ · polymers, in particular, are well suited...

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