Review of Adaptive Programmable Materials and Their

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Review

A Review of Adaptive Programmable Materials and Their Bio-applications Xiaoshan Fan, Jing Yang Chung, Yong Xiang Lim, Zibiao Li, and Xian Jun Loh ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.6b09110 • Publication Date (Web): 23 Nov 2016 Downloaded from http://pubs.acs.org on November 24, 2016

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ACS Applied Materials & Interfaces

A Review of Adaptive Programmable Materials and Their Bio-applications

Xiaoshan Fan1, Jing Yang Chung3, Yong Xiang Lim3, Zibiao Li2*, Xian Jun Loh2,3,4*

1

School of Chemistry and Chemical Engineering, Henan Normal University, China

2

Institute of Materials Research and Engineering (IMRE), A*STAR, 2 Fusionopolis Way, Innovis,

#08-03, 138634 Singapore 3

Department of Materials Science and Engineering, National University of Singapore, 9

Engineering Drive 1, Singapore 117576, Singapore 4

Singapore Eye Research Institute, 11 Third Hospital Avenue, Singapore 168751, Singapore

Correspondence: X. J. Loh ([email protected]); Z. Li ([email protected]) Tel: +6565131612

Keywords: Adaptive programmable, Stressed bilayer, Patterned film, Shape memory, Drug delivery,

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Abstract: Adaptive programmable materials have attracted increasing attention due to their high-functionality, autonomous behaviour, encapsulation and site specific confinement capabilities in various applications. Compared to conventional materials, adaptive programmable materials possess unique single materials architecture which can maintain respond and change their shapes and dimensions when they are subjected to surrounding environment changes such as alternation in temperature, pH and ionic strength. In this review, the most recent advances in the design strategies of adaptive programmable materials are presented with respect to different types of architectural polymers, including stimuli responsive polymers and shape memory polymers. The diverse functions of these sophisticated materials and their significance in therapeutic agent delivery systems are also summarized in this review. Finally, the challenges for facile fabrication of these materials and future prospective are also discussed.

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1. Introduction Polymers have been widely used as biomaterials. Common industrial polymers used in manufacturing were repurposed for the first generation of biomaterials, like the use of Dacron for vascular grafts and poly(methyl methacrylate) (PMMA) for contact lenses. Today, polymers have been used in medical devices like stents, heart valves, pacemakers and hip implants1. The wide range of biomedical applications for polymers stem from the variety of properties that different polymers possess. Many biomaterials are non-toxic to the human body, such as naturally derived cellulose, chitosan and chitin or consisting of biocompatible monomers such as glycolic and lactic acid. Importantly in the design of third generation, biodegradable medical devices, some polymers like poly(lactic acid) (PLA) and poly(glycolic acid) (PGA) can degrade in the human body.2-6 Also, many polymers can be easily moulded and shaped using techniques like injection moulding. Due to these useful properties, polymers are an essential class of material for engineers to consider as a biomaterial in designing new medical devices. In recent years, much research has been done on the possible uses of polymers in drug delivery systems. There has been growing recognition that having controlled, sustained and site-specific delivery of drugs to disease sites leads to better patient outcomes, compared to the current conventional drug delivery methods that can cause sharp increases in drug concentrations to potentially toxic levels.7-11 In particular, biodegradable polymers have been considered as delivery vehicles, acting as a reservoir to hold and protect bioactive drug molecules until processes within the human body like hydrolysis or enzymatic attack cause the degradation of the protective polymer shell, leading to the release of the bioactive molecule.12-17 Such biodegradable polymers have been used to create microspheres, which have been studied extensively both in vitro and in vivo for their degradation profiles and drug delivery18. In order to effectively deliver the bioactive molecule, it must first be encapsulated within the biodegradable polymer shell. This encapsulation can act to enhance solubility, increase site-specificity, prevent premature degradation, permeate barriers like cell walls, and reduce dosages to limit side effects of the drug19-21. While the aforementioned 2 ACS Paragon Plus Environment


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methods of encapsulating drugs may address some of these characteristics, there are limitations in regards to including multiple attributes within the same manufacturing process. One main limitation is that, even though methods like double emulsion can achieve reasonable homogeneity, manufacturability and encapsulation, the structures produced are primarily limited to a spherical shape, which as mentioned earlier has huge implications of its effectiveness due to the effect of shape and size of drug delivery vehicles in the circulation and retention of the drugs in the human body. Such spherical vehicles are also unable to provide directional release of drugs. In addition, there has been difficulty in controlling the thickness and porosity of the membrane wall for gel capsules22-24. 1.1 Programmable Adaptive Polymers The concept of programmable materials refers to a self-assembly type of process whereby flat and planar polymer structures spontaneously change their shape and dimensions. These programmable adaptive polymers are within a class of stimuli-responsive materials known as smart polymers, and the usage of these materials is a new approach being considered to address the limitations of traditional drug encapsulation. The driving force for programmable adaptive materials arises from the special characteristics of these materials, whereby a sharp discontinuous change in their properties is exhibited upon a small or modest change in their physical environment such as through variations in temperature, pH, chemicals, electrical signals, differential stress or magnetic fields25. These changes then allow the material to spontaneously fold into 3D structures of various shapes. In the case of drug delivery systems, programmable adaptive polymers, rely on these large changes to bring about their controlled release of drugs into the body. Presently there are three methods for the design of programmable adaptive polymers. The first approach utilizes shape-memory polymers with directional anisotropic behaviour. Similar to shape-memory metallic alloys, these polymers could be crafted into a temporary shape based on its applications, and recover its original shape upon being heated. The second approach utilizes patterned films consisting of two or more polymers, with one being active and the others passive. The active polymer region acts as hinges to 3 ACS Paragon Plus Environment

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induce folding. The third approach incorporates two or more polymers in the same system with different swelling properties and is commonly called the differentially stressed method. Due to the different swelling rate in response to a signal, the different parts of the system shrink/expand at different rates allowing folding. All the methods are shown in Figure 1.1.

Figure 1.1. Methods of designing programmable adaptive functionality into polymers. The transition from the unfolded shape to the folded shape is depicted for (a) the shape-memory polymer, (b) the patterned film polymer system, and (c) the differentially stressed polymer system. Thermodynamically, the swelling or contraction of polymers systems is made possible through the changes in the conformation of polymer chains to reduce the free energy of the system by either ‘contraction’ (the bunching up of polymer molecules) or ‘swelling’ (stretching of the polymer chain). In solutions, the polymer coil can be represented as an equivalent hydrodynamic sphere that is, in effect, impenetrable to solvent molecules. The size of this sphere, commonly represented by the radius of gyration, thus determines how much ‘contracted’ the coil is. The Gibbs energy of mixing, , is determined by the enthalpy of mixing, , and the entropy of mixing, , .

(1.1)

In the case of long chained polymer-containing mixtures, the basis for the theory of mixing has been laid down in the form of the Flory-Huggins solution theory26, which takes into account the dissimilar molecular sizes and considers the polymer-polymer, polymer-solvent and solvent-solvent 4 ACS Paragon Plus Environment


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interactions in adapting the usual expression for the entropy of mixing. Although there are assumptions made for this theory, namely, that the volume change upon mixing is zero and that molecules in the chain have an equal probability to occupy all lattice sites, these are negligible for most purposes27. These assumptions then allow the derivation of the Gibbs free energy of mixing of a polymer solution,

ΔG RT[(1 Φ )ln(1 Φ ) +

where

lnΦ + Φ (1 Φ )χ ,

(1.2)

is the interaction parameter taking into account the interactions between polymer and

solvent molecules while also encompassing the entropic contribution resulting from the mixing of different species, Φ is the composition variable for polymer species 1, which gives the segmental molar fraction of the polymer and N is the number of segments forming the polymer chain. We could thus see that the conformational behaviour is highly dependent on χ. If interactions between polymer and solvent molecules are energetically favourable, the hydrodynamic sphere would favour swelling. On the other hand, if the polymer and solvent molecules interactions are poor such that polymer-polymer self-interactions is preferred, the hydrodynamic sphere would contract.

If the solvent happens to be another polymer species 2, the equation would then be,

ΔG RT( lnΦ + ! lnΦ" + Φ Φ" χ) .

(1.3)

!

Fundamentally, equation 1.3 provides the basis for our understanding of the origins of folding. Chain folding is not limited to polymer chain structures and can also occur in cylinders, spirals, or corrugated sheets. For a cross-linked polymer system, a stable network and a reversible switching transition between its folded and non-folded states are two prerequisites for programmability and adaptability28. The stable network determines the original shape of the system while the lock in the


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network is responsible for its temporary shape. The mixing for such systems can be described by the Flory-Rehner equation29,

*

[#$(1 % ) + %" + %" " & ' $ (%" ) ! +,

(1.4)

"

Where %" and %" are the volume fraction of the solvent and polymer, n is the number of network chain segments bounded on each ends of the crosslink. This equation considers the changes in the enthalpy of mixing and changes in the entropy both from mixing and the reduction of possible chain conformation from swelling. If the hydrogel is non-degradable, the size of the system needs to be below the renal threshold to ensure that accumulation in the body does not occur.30-34 Such folding can be observed in naturally occurring bio-polymers with proteins and nucleic acids forming complex 3D structures and is an integral feature for the workings of a cell. These are generally mediated by relatively weak interactions such as through hydrogen bonding. In vesicle formation, spontaneous curving and folding of lipid bilayers, thin molecular films, into spherical shapes have been observed35. The programmable adaptive process can allow greater control over the shapes and sizes of drug delivery vehicles, making it possible to create non-spherical polymer shells. In addition, programmable adaptive techniques, in conjunction with methods of creating patterned thin films like electrospinning and 3D printing can lead to greater control of the wall thickness, porosity, surface patterns and composition of polymeric containers36, all important factors that decide the effectiveness of the drug delivery vehicle. Currently, programmable adaptive polymeric containers have been created at the mm to 100 nm scale and research efforts are being directed to achieving smaller containers for drug delivery. Design principles for programmable adaptive polymers include the control of steric interactions and molecular rigidity. While the mechanisms and techniques used in programmable adaptive materials at different size scales may differ, the geometric principles for programmable adaptive materials appear to hold through across different sizes37. Such programmable adaptive polymeric containers



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have been used to successfully encapsulate a variety of bioactive molecules, ranging from proteins, peptides, fungi, bacteria and cells38. There are two main types of programmable adaptive polymer structures. The first consists of programmable adaptive polymer chains, which is analogous with the self-folded biomolecules like proteins that can be observed in nature. The second consists of thin films of polymers, which is most commonly used for drug encapsulation and delivery. Thin films of polymers can be deposited onto a surface using dip casting, spin casting or other precipitation techniques from a polymer solution. The surface characteristics, ligand adhesion regions and porosity of these thin films can then be controlled by micro and nanofabrication techniques like electron beam lithography and focused ion beam lithography that pattern the thin film. Different shapes for the polymer films can also be obtained through cutting39, 40, using microwell-like substrates41and photolithography42. Each of these methods have their own advantages. Cutting is by far the simplest of the methods and can be used for all types of cross-linkable polymers, although is best used to form simple shapes only. Using microwell-like substrates allow the formation of more complicated shapes such as stars. Large scale production of polymer films can be made through photolithography. With the suitable surface characteristics and shape, various stimuli can then be applied to the polymer film to induce programmability and adaptivity through various mechanisms. There are two fundamentally different effects used to generate shape changes in polymeric systems, namely, a) differential swelling (which is a result of solvent quality); and b) shape memory (which is an effect of chain mobility. For the first, it is the result of the change of the solvent quality.

Changes at the

temperature, pH value and the solvents change the quality of the solvent, resulting in a change in the conformation of the polymer chains. This result macroscopically appears as swelling. For shape memory, it is normally accepted as an effect seen in solid polymer systems that a frozen into a nonequilibrium macroscopic shape. The temperature used to induce shape memory affects chain mobility – increasing the temperature allows more rotational motion about the polymer backbone, and allows the chains to adopt the lowest energy conformation. In this case, the solvent quality 7 ACS Paragon Plus Environment

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needn't brought to consideration because the polymer is not solvated. This paper will give an overview of these stimuli before examining the different mechanisms as utilized in thin films programmable and adaptive polymers. 2. Stimuli 2.1 Thermal Responsive Polymers Due to the presence of the temperature term within the thermodynamics of mixing equation, temperature is the most straightforward stimuli to describe.43-45 Here we will also introduce two terms related to miscibility and temperature; the Lower Critical Solution Temperature (LCST) and Upper Critical Solution Temperature (UCST). These two temperature points are respectively, the lowest and highest temperature of the binodal/spinodal curves by which the components of a mixture are entirely miscible and the solution is homogenous. The other two types of phase behaviour are total immiscibility and partial miscibility which lies above or below the LCST/UCST. While the upper and lower temperature curves appear similar, LCST is entropically driven while UCST is enthalpically driven46, and these behaviour result from the polymer and solvent interactions taking place. Polymer systems with a LCST will display negative free energy when the temperature increases resulting in water-polymer association being unfavourable as shown in Figure 2.1.46 This effect is known as the hydrophobic effect. For low molecular weight polymers, a higher temperature would increase the ΔS term in the thermodynamic expression more than other contribution, leading to higher miscibility. However, for high molecular weight polymers, the

ΔS term is far less important and other entropy contributions and temperature dependantΔH values will contribute more and lead to the reverse behaviour. Thus, thermoresponsive polymers can, in general, display either of these critical temperature points or both with a miscibility gap in their temperature-composition diagram. This property then allows the conformational changes to the coil as described above and the LCST/UCST properties can be tweaked by altering the degree of polymerization as well as polydispersity index. The sol-gel transition temperature of these polymers can be experimentally determined using various techniques such as spectroscopy47, differential 8 ACS Paragon Plus Environment


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scanning calorimetry48 or rheology49. For programmable adaptive polymeric systems utilizing cross-linking for drug delivery, two approaches can be made. One approach is that the hydrogel cross-linked nodes can be formed in situ after being exposed to the body’s environment. The other approach would be to form the cross-linked network before implantation in the body. Regardless of the approach, the control of the release of drugs from the folded/unfolded polymer system would be achieved upon heating to body temperature, which induces the change in the physical properties of the polymer system. While polymer-solvent mixtures tend to exhibit an UCST, polymer-polymer mixtures or polymer blends would exhibit a LCST. Common examples of LCST smart polymers are poly(N-isopropylacrylamide) (PNIPAM) with a critical temperature of 32 °C50 and is one of the most widely studied polymer in the biomedical field due to this temperature being very close to body temperature, poly(N,N-diethylacrylamide) (PDEAAm) with a critical temperature range of 25°C to 32°C51, poly(N-vinlycaprolactam) (PVCL) with a critical temperature range of 25°C to 35°C52, poly(2-(dimethylamino) ethyl methacrylate) (PDMAEMA) with a critical temperature of around 50°C and poly(ethylene glycol) (PEG) / poly(ethylene oxide) (PEO) with a critical temperature of 85°C53. PNIPAM-based systems display a sharp phase transition where a flexible coil structure exist below the LCST. Above the LCST, the structure becomes hydrophobic and the chains collapse, causing aggregation and leads to folding of the overall structure in programmable adaptive systems. The phase transition of PNIPAM systems is influenced by the presence of salts, pH and other polymer species as the copolymerization of PNIPAM had been widely studied50. However, an issue with these systems is on biodegradability. Many homopolymer and copolymers of NIPAAm are not biodegradable unless purposefully mixed with biodegradable polymers in the chain. The other important group of thermoresponsive systems comes from the PEG/poly(propylene oxide) (PPO)-based polymers. Triblock of PEO-PPO-PEO, which is also known as Poloxamers, display thermoreponsive behaviours over a wide range of temperatures and can be tweaked by adjusting the composition, molecular weight and the concentration of the species54. The amphiphilic behaviour of this system is made possible by the opposing hydrophilic properties of ethylene oxide 9 ACS Paragon Plus Environment

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and the hydrophobic properties of propylene oxide. Gelation is brought about by changes in the micellar properties. An issue with this copolymer is its low mechanical properties and poor viscosity55, which could cause the release of drugs to proceed too quickly. 2.2 pH Responsive Polymers In the human body which has a pH of close to 7.4, variations on pH can occur in various body site, for example, the blood vessels or the gastro-intestinal tract. A large change in pH occurs moving from the acidic environment in the stomach (pH = 2) to the more basic environment in the intestinal tract (pH = 5-8). Cancerous portion in the body as well as inflamed tissue also display a different pH as in circulation56. For instance, chronic wounds could have a wide range of pH from 5.4 to 7.4 and cancer tissue is known to be extracellularly acidic in nature due to the low oxygen partial pressure and high metabolism rate57. These places act as the base for the release of drugs for programmable and adaptive drug delivery systems. In general, pH-responsive cross-linked systems composed of polymeric backbones containing ionic pendant groups and in aqueous environments of the correct ionic strength and pH, the pendant groups ionize and cause fixed charges to appear on the polymer backbone. This then leads to the generation of electrostatic repulsive or attractive forces which induce the swelling or deswelling of the pH responsive hydrogel which causes folding (Figure 2.1).58 Thus, the polymers forming up the hydrogel are essentially polyelectrolytes. Here we ought to introduce a term known as the acid dissociation constant or ./ which measures the strength of an acid in solution. The larger the value of ./ , the larger the extent of dissociation of the acid,

./

[01 &[2 3 & [20&

.

(2.1)

More commonly used is the logarithmic constant of the acid dissociation constant, 4./ , which allows comparison with pH. For anionic pH responsive hydrogels, its pendant groups will ionize when the subjected environment pH is above its 4./ . If the pH is below its 4./ , the opposite occurs and the pendant groups deionize. The reverse is true for the case of cationic hydrogels and swelling occurs at lower pH. Continuous ionization of the pendant groups in polyelectrolyte may be 10 ACS Paragon Plus Environment


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hindered due to the effects of the adjacent ionized groups, which will exert an electrostatic force. This results in the apparent dissociation constant ./ to be smaller than that of its monoacid or monobase.

Figure 2.1. Depiction of the swelling/gelling transition upon changes in the pH for (a) an anionic hydrogel and (b) a cationic hydrogel. Reprint from Elsevier with permission.58 All pH sensitive polymers contain either pendant acidic (e.g. carboxylic acid, sulfonic acid) or basic (ammonium) groups which can take in or give out protons during changes in the pH. Widely studied pH responsive ionic polymers include poly(acrylic acid) (PAA) and poly(methacrylic acid) (PMAA) which ionizes at high pH, poly(diethylaminoethyl methacrylate) (PDEAEMA) and poly(acrylamide) (PAAm), poly(dimethylaminoethyl methacrylate) (PDMAEMA) which ionizes at low pH59. Apart from synthetic polymers, a number of natural polymers also display pH-responsive behavior such as albumin and gelatin which occurs through hydrogen bonding at suitable pH and temperature. At their isoelectric point, these proteins contain the lowest amount of surface charges and at pH away from their isoelectric point, display swelling due to an increase in surface net charge and increased electrostatic repulsive force. There are a number of factors beyond the 4./ of the hydrogel which influence the degree of swelling and the subsequent folding of programmable adaptive systems. As the swelling is due to the electrostatic repulsion between the pendant groups on the polymer chain, anything which increase or reduces the electrostatic repulsion would have an effect. The nature and concentration of the buffering agent affects the kinetics of the gelling, for instance, the rate of swelling of poly(methyl

methacrylate-co-N,N-dimethylaminoethyl

methacrylate)(PMMA-co-PDMAEMA)

increases when buffered by weak organic acids60, 61. Monovalent anions in general are affected to a 11 ACS Paragon Plus Environment

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greater degree compared to multivalent anions in the influence of buffers. Also, the greater the 4./ of the acidic buffer, the greater the folding rate of the polymer system will be. The physical structure of the pH responsive hydrogel is also a key factor in the extent of swelling62, for example, the degree of cross-linking, ratio of pendant groups, molecular weight, etc. 2.3 Solvent Responsive Polymers In general, solvent responsive systems contains neutrally-charged water-swelling polymers which folds when immersed in an aqueous medium. The mechanism for folding is inherently due to the interactions between polymers and the solvent and was described in depth through Equations (1.2 – 1.4), where the degree of swelling was shown to depend on the interaction parameter χ. This interaction parameter can be further expanded out63,

χ

67 897 :9; < =>

!

,

(2.2)

where?@ and ?A are the solubility parameters of the solvent and polymer species, '@ is the molar volume of the solvent and B is the molar gas constant. The value for the solubility parameter of a species depends on its chemical structure64and the smaller the difference between the solubility parameter of the solvent and polymer, the more like dissolution will be. Examples of solvent responsive polymer systems are polyvinyl alcohol (PVA)-chitosan, chitosan-poly(ethylene glycol)65, and polydimethylsiloxane (PDMS)–polyurethane/2-hydroxyethyl methacrylate which swells upon contact with water, and poly(glycidyl methacrylate) (PGMA) – gold66 which swells upon contact with methanol. 2.4 Other Stimuli Responsive Polymers Beyond temperature, pH and solvent stimuli, there are other stimuli, albeit less common, which can also trigger the folding of the polymer system. The presence of an electric field can induce an electrostatic repulsion to polyelectrolyte hydrogels and cause swelling or shrinking. This is in essence, triggering what would normally be pH sensitive polymer systems to fold or unfold without needing an alteration in the pH. For instance, electric current applied in a pulsatile manner can 12 ACS Paragon Plus Environment


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allow the release of drugs encapsulated in poly(2-acrylamido-2-methylpropane sulfonic acid -co-nbutylmethacrylate) (PAMPS-co-n-BMA）as well as in polyacrylamide67. A programmable adaptive system of polydimethylsiloxane – cardiomyocyte, which adopted three dimensional conformation as an electric field was applied, was shown to display gripping and walking motions68. Although drug delivery using these electro-responsive polymer systems seems novel as the release rates can easily be controlled, works on this method are still in their infancy. Problems faced by this method include the requirement of a constant external voltage source, which may not be the most convenient of methods, as well as needing to work in the absence of electrolytes, which is not achievable under most physiological conditions. Other stimuli which could be considered are light, glucose and enzyme, to name a few. Light sensitive programmable adaptive-systems have the advantage of being able to have the stimuli delivered in specific amounts instantaneously, and with high accuracy69. Glucose sensing systems can work as a modulating insulin deliver67 while enzyme-based strategies allows no prior knowledge of the specific position of the sites which requires drug release such as tumours70. 3. Differentially Stressed Thin Films 3.1 Simple Rectangular Bilayers Bilayers are a straightforward way to cause bending of folding of a material. The use of bilayers to create mechanical movement and folding is not new and has been used for centuries in the bimetallic strip. A bilayer polymeric film typically consists of an active polymer layer that is deposited onto a substrate, or passive layer. Each layer will have different mechanical or swelling properties. Upon application of a stimulus that causes a change in volume, differential stress occurs as the two layers have different mechanical properties. Three possible scenarios will then happen depending on the properties of the thin film.71 Firstly, if the passive layer is highly resistant to deformation, the active layer will either wrinkle or crease as shown in Figure 3.1(a) and Figure 3.1(b) below. On the other hand, if the passive layer is soft and deformable, the folding or bending of the entire structure becomes possible as seen in Figure 3.1. 13 ACS Paragon Plus Environment

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Figure 3.1. Possible scenarios in the deformation of polymeric thin films (a) wrinkling, (b) creasing and (c) folding. Reprint from the Royal Society of Chemistry with permission.71 While the structures created by folding is different compared to that created by wrinkling and creasing, the underlying mechanisms are related. Cendula et al. found that for compressively strained thin films, the magnitude of the strain gradient in the active film was responsible for determining whether wrinkling/creasing or folding occurred. At lower strain gradients, the wrinkling/creasing of the films was observed. As strain gradients increase, however, a phase transition from wrinkling to bending was observed72. At the most basic level, the relationship between the stress, thickness of a polymeric bilayer and the resultant curvature from bending can be approximated by the Timoshenko equation for bending in metallic beams73,

1 6(ε" ε )(1 + m)" p h (3(1 + m)" + (1 + mn)(M " +

J

)+

,

(3.1)

where n E /E" and m a /a" , E and E" are the modulus of elasticities for the two polymer layers, a and a" are the thickness of the layers, h is the total thickness of the film and p is the radius of the curvature of the rolled up polymer layers. The Timoshenko equation predicts that the radius of the curvature is inversely proportionate to the stress in the film. It also predicts that the radius of curvature will increase initially with m but will decrease past a certain point. However, the equation cannot predict the direction of bending. In addition, the Timoshenko equation was developed for metallic materials, which have relatively small changes in volume. In contrast, polymers exhibit much larger volume changes, and thus require separate study to characterise their folding properties. A recent study by Stoychev et al.74 focused on rectangular hydrogel based 14 ACS Paragon Plus Environment


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polymer bilayers. They found that the ratio of the width and length of the rectangular bilayers was important in determining the direction of folding/rolling and their results are shown in Figure 3.2 below. Large width to length or length to width ratios led to folding along the long side of the rectangle. A large actuation (i.e both length and width are much greater than the rolled circumference) leads to folding from all sides while a small actuation (i.e. both length and width are comparable to the rolled circumference) leads to diagonal folding instead.

Figure 3.2. Rolling directions of rectangular polymeric bilayers. Reprinted from Elsevier with permission.75 In a similar vein, Morimoto & Ashida75 have examined the folding behaviours of a bilayer consisting of a swelling hydrogel and an elastomer layer. In the absence of external forces, they found that the aspect ratio L/H influences on the semi-angle while the curvature does not, and the normalized crosslink density has an effect on both the bending semi-angle and curvature. They also found that by quantifying the stress distributions at T = 300 K, multiple neutral axes can exist in the bilayer under certain conditions. They also indicated that the folding shape of the bilayer can be controlled by tuning both the ratio of shear modulus between the two layers and the aspect ratio L/H. 3.2 Bilayers with Lateral Swelling Differentials Similarly to bilayers with differential changes in volume, two polymer thin films attached laterally will also fold upon differential changes in volume (Figure 3.3).76 Compared to the normal bilayers, these laterally attached thin films will have both mean and gaussian curvature. This method of inducing folding of a 3D shape has also been observed in nature in the folding of thin sheets of 15 ACS Paragon Plus Environment

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cells77. Kim et al.76 investigated thermal responsive hydrogels based on this architecture. A photolithographic technique was used on a single sheet of P(NIPAm-co-BPAm-co-AAc-coRhBMA) to create differences in the extent of crosslinking in each half of the sheet. Kim et al. found that the polymer film would roll into a cylinder like structure with a neck at the interface between the two films at 22 °C. The size and morphology of the neck was dependant on factors like temperature and the thickness ratio between the two layers. Upon raising the temperature to 50 °C, the difference in swelling between the more cross-linked half and the less cross-linked half would be minimized, and the structure will unfold into a flat sheet.

Figure 3.3. Different folding in bilayers with (a) vertical swelling differential and (b) lateral swelling differential. Reprinted from the Royal Society of Chemistry with permission.76 3.3 Fabrication of Bilayers The manufacture of polymeric bilayers can be done in a few ways. A substrate may be used to create a base upon which the bilayers can be deposited. This substrate could be glass or another polymer. Spinning of the polymer onto the substrates can then be done to create layers upon the substrate. If metallic layers are to be used for the passive layer, techniques like sputtering can be used to deposit the metallic layer. The bilayers may be chemically linked depending on the system, with methods like UV exposure used to create chemical crosslinks between polymer layers. The rectangular shapes of the bilayers can then be mechanically cut. 3.4 Temperature Responsive Bilayer Systems Temperature responsive polymeric bilayers are bilayers that fold in response to changes in temperature. A few such systems have been developed towards drug delivery. Stroganov et al.78 have developed simple temperature sensitive biodegradable self-rolling cylinders using gelatin as 16 ACS Paragon Plus Environment


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the active polymer and polycaprolactone (PCL)/PHF as the hydrophobic substrate. These polymers were chosen due to their biodegradability and biocompatibility. The temperature dependant bending is driven by the swelling behaviour of gelatin. Gelatin is capable of forming a hydrogel and the solgel transition is governed by temperature. The swelling of the hydrogel in comparison to the passive layer causes folding. Stroganov et al. studied two cases: a non-crosslinked bilayer of gelatin/PCL and a crosslinked bilayer of gelatin/PHF-Q. In the non-crosslinked gelatin/PCL bilayer, the bilayer was observed to fold into a cylinder at 22 °C within 1500 s. Heating of the bilayer to 37 °C above the sol-gel transition temperature caused the gelatin layer to dissolve, leaving behind an unfolded PCL layer. For the cross-linked system, the gelatin layer is larger than the PHF-Q layer, ensuring that after UV crosslinking, sections of the gelatin remain uncross-linked. Upon raising the temperature to 37 °C, the non-crosslinked sections of the gelatin dissolved. The remaining crosslinked gelatin also swelled, but was unable to dissolve into solution due to cross-links with the PHF-Q layer and caused the bilayer to fold due to the swelling of the cross-linked gelatin. The gelatin/PHF-Q system was also demonstrated to be successful in encapsulating the neural stem cells of mice. The stem cells were added to the bilayer at 22 °C and attached to both hydrophobic and hydrophilic portions of the bilayer. Upon increasing the temperature to 37 °C, the bilayer folded into a cylinder and encapsulated the mice cells. The cells in the tubes were found to be stable after seven days. This demonstrated the possibility of these bilayers for drug delivery with the gelatin/PCL unfolding and releasing its payload at 37 °C, or at body temperature, while the gelatin/PHF-Q system would continue to hold its payload at 37 °C in the human body and slowly degrade, creating a more timed release. Other thermal responsive systems that do not rely on hydrogel swelling have also been developed. Simpson et al.40 designed a residual stress based thermal responsive folding polymer bilayer system. PDMS is used as an active layer and is deposited onto a PAA sacrificial layer, following which a metallic layer of Ti or Au is deposited onto the PDMS layer using e beam evaporation. Upon deposition onto a PAA sacrificial layer, a residual strain develops between the PDMS-Au 17 ACS Paragon Plus Environment

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bilayer due to lattice misfit strain between the layers. This bilayer then rolls up into a cylinder upon release from the PAA layer as the PDMS layer shrinks to relieve the interfacial stress. Upon heating, the bilayer unrolls into a flat 2 dimensional sheet due to the difference in thermal expansion coefficients between the two layers. This process is reversible due to the elastomeric properties of the PDMS. The control of the radius of the cylinders can be achieved by the controlling the curing temperature of the PDMS. Two competing effects were identified by Simpson et al. Increasing the curing temperature causes a higher residual strain to develop within the PDMS layer, thus leading to more bending and smaller radius of the cylinders upon unfolding. However, the elastic modulus of the PDMS layer will also increase as temperature increases, leading to a stiffer layer that is more resistant to bending. Simpson et al. thus reported an optimum temperatures of around 110 – 130 °C for creating small microparticles with diameters around 160 microns. Changing of the passive layer also affects the microparticle size as it is highly dependent on the mismatch in mechanical properties between the PDMS and passive layer. Simpson et al. investigated the use of SiC as a passive layer instead of Au and found that SiC layers were not less effective than metallic layers as thick SiC layers were required to cause the same residual stress development in the PMDS-passive layer interface. The rolling of the bilayer was also not as uniform in the PDMS-SiC bilayers compared to the PDMS-Au bilayers. The potential of the PDMS-Au system in encapsulating and releasing polymers was also examined by Simpson et al. through the capturing and release of PEG. The PDMS-Au bilayer was successful in capturing PEG at room temperature and releasing PEG upon unravelling at higher temperatures. However, it was noted that this encapsulation would be limited to hydrophilic molecules as the Au layer is facing inwards with the PDMS layer facing outwards, leading to hydrophobic molecules preferring to adhere to the outer walls. It is thus possible for proteins with hydrophilic domains to be encapsulated within this bilayer and released inside at the elevated temperatures of the human body. Zakharchenko et al.79 used a similar approach with a thermally responsive bi-layer to create a rolled up cylinder to encapsulate cells. The film utilized a stiff PCL passive layer and a thermal 18 ACS Paragon Plus Environment


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responsive PNIPAM hydrogel layer to create a bilayer. The bilayer was observed to roll at temperatures below 28 °C and unroll at temperatures above that rapidly in about 1-3 seconds. The encapsulation and release capabilities of the tubes were tested with 10µm SiO2microparticles. The SiO2 particles were trapped upon folding of the bilayers at lower temperatures and released upon the unfolding of the bilayers at elevated temperatures, highlighting the possibility of encapsulating cells that are within the same size range. The possibility of manipulating the cylinders using magnetic fields was also examined by adding Fe3O4 microparticles. The microtubes were then found to move and change directions towards magnets, thus demonstrating the possibility of directing the bilayer microtubes towards specific portions of the human body through magnetic fields.

3.5 pH Responsive Bilayer Systems Generally, pH responsive bilayer systems rely on a layer being a weak polyelectrolyte which acts as the active layer. Ding et al. utilized parylene and a dry hydrogel to form a bilayer microstructure where the hydrogel layer shrinks at low pH80. Parylene or poly(p-xylylene) is a commonly used as a biomaterial for its extreme stability to chemical attack, low dissipation factor and high mechanical properties. In general, parylene resists environmental changes very well which makes it suitable to work as the passive layer in the structure. This bilayer system is simplistic in nature due to the strong adhesion between parylene and the dry hydrogel layer without needing modification. With an increase in the environmental pH, the hydrogel switches from a contracted structure to a swelled structure due to ion diffusion and osmotic pressure. Parylene does not undergo conformational changes and as such, the structure folds with the extent of the curvature being influenced by the swelling ratio of the hydrogel, as illustrated in Figure 3.4.


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Figure 3.4. Representation of a hydrogel – parylene microstructure transitioning from its (a) unfolded state to its (b) folded state as the pH increases. Reprinted from with permission.80

Ding et al. observed that high crosslinked hydrogel is necessary to control the extent of the folding of the bilayer fingers. To vary the amount of folding, it was also reported that using a thicker hydrogel layer would increase the amount of deflection. The hydrogel was observed to swell not only in the direction of the shank but perpendicular as well. Self-rolling films have also been studied and described. Luchnikov et al. showed that polystyrene-poly(4-vinyl pyridine) bilayer could roll at low pH due to the swelling of poly(4-vinyl pyridine)39. As the pH decreases, the rate of folding

increases.

The

same

author

also

tested

polystyrene-poly(4-vinyl

pyridine)–

polydimethylsiloxane trilayer which also demonstrated rolling at low pH due to the conformational changes of poly(4-vinyl pyridine). In another study, He et al. utilized pH sensitive poly(methacrylic acid) – poly-(2-hydroxyethyl methacrylate) (PMAA-PHEMA) as well as poly(methacrylic acid) – (PMAA)/poly(ethylene glycol dimethylacrylate) (PMAA-PEGDMA) which folds upon immersion in biological fluids81. A relatively large millimetre-sized pH responsive system was demonstrated by Gracias et al. who used poly(ethylene glycol)- poly(N-isopropylacrylamide-acrylic acid) (PEGP(NIPAAm-AAc)) bilayers to induce snapping upon changes in the pH82. Thin film based pH responsive folding systems have also been developed. Rajagopal et al. managed to control the trigger of the folding, self-assembly, and subsequent hydrogelation of amphiphilic β-hairpin peptides through pH83. When immersed into solutions of high pH above 9, the peptide remains 20 ACS Paragon Plus Environment


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unfolded. The unfolded nature is a result of the electrostatic repulsion which would arise due to the vast number of protonated lysine side chains, which would have to occupy the same face of the molecule if folding were to occur. However, it was also noted by Rajagopal et al. that higher temperatures would make folding thermodynamically favourable due to the dissolution of the hydrophobic valines, which would allow the peptide to fold and self-assemble. The equilibrium of the peptide folding and the self-assembly of the system are linked. The rate of the peptide folding, and subsequent self-assemble is related to the number of crosslinks that develops throughout the gelation process, which directly influence the mechanical properties of the gel. Thus, slow kinetic rates for folding lead to less rigid gels while fast kinetic rates of folding lead to more rigid gels. Rajagopal et al. also tested the peptide by changing its net charge through the replacement of lysine with threonine and isoleucine, which causes the peptide to fold at pH lower than that of lysine. The study found that there is a relation between the net charge of the peptide and its ability to fold, as well as the position of the substitution within the peptide sequence indicating that peptides can be manually designed to fold at specific pH.

3.6 Solvent Responsive Bilayer Systems Solvent responsive programmable adaptive films are films which fold upon immersion in solutions, be it in aqueous medium or otherwise. For instance partially biodegradable poly(vinyl alcohol)– chitosan bilayer systems as well as chitosan –poly(PEGMA-co-PEGDMA) poly(poly(ethylene glycol) methylacrylate-poly(ethylene glycol) dimethylacrylate) (P(PEGMA-PEGDMA)) have been fabricated and demonstrated by Lee et al. which folds upon contact of an aqueous medium due to the swelling of polyvinyl alcohol and P(PEGMA-co-PEGDMA)65. These microstructures were made through the usage of microwell-like substrates where chitosan was crosslinked by glutaraldehyde while P(PEGMA-co-PEGDMA) was crosslinked through the exposure to UV light successively. The fabrication process involves microtransfer moulding, discontinuous dewetting and sacrificial layer techniques. Because the different layers have different swelling ratios in 21 ACS Paragon Plus Environment

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aqueous mediums, the bilayers self-fold from 2D structures to 3D structures. The adhesion between the two different layers is achieved through the diffusion of PEGDMA and PEGMA monomer molecules into the chitosan layer. These monomers then polymerize within the chitosan which forms a strong interface to allow the two layers to adhere to one another. The degree of folding was found by Lee et al. to depend on the PEGMA to PEGDMA ratio where a higher ratio led to a greater deflection of the microstrip. For the development of other aqueous folding solvent systems, Jeong and Jang et al. utilized the water swelling properties of polydimethylsiloxane–polyurethane-2-hydroxyethy methacrylate complex bilayers to induce folding and found that the structure were was able to form different 3D objects such as tubes, cubes, pyramids and helixes84. When these systems come into contact with water, they immediately fold due to the aqueous medium hampering the loading of cells. Interestingly, non-aqueous solvents can also be used to trigger the folding of polymer bilayer systems. Kelby

et al. reported methanol as a stimuli for a system based on poly(glycidyl

methacrylate) (PGMA) brush layers grafted to gold patterned films66. Although polymer-metal bilayers have been studied as bioactuators, those systems utilized bulk polymers rather than brush polymers as utilized by Huck et al. The advantage of polymer brushes is that interchain interactions which can be attractive or repulsive can give rise to swelling or contracting. For this system, the characteristics of the polymer brush and the Au thickness will influence the amount of folding. Through ellipsometry measurements, PGMA brushes were found to have a greater swelling factor in methanol than water. Crosslinked PGMA compared to non crosslinked PGMA were also found to have a greater swelling ratio. 3.7 Non Stimulus Triggered Bilayer Systems All the examples described thus far utilized a stimulus as a trigger for the folding or unfolding process. However, a bilayer system which self-assembles through folding and releases drugs while keeping its shape constant is possible as well. Baek et al.85 reported on such a programmable adaptive hydrogel bilayer system (Figure 3.5). The bilayer consisted of PEGDA layers with each 22 ACS Paragon Plus Environment


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layer having a different crosslink density. The difference in cross-link density leads to differential swelling when placed in an aqueous medium. Exposure to UV light was used to control the density of cross-links in each PEGDA layer. The increase in molecular weight from cross linking in the top PEGDA layer increased its expansion ratio. This differential swelling of the layers led to folding into cylindrical tubes. The radiuses of the cylinders formed were found to depend on the concentration of the polymer in the inner polymer layer. Baek et al. also tested the released properties of the cylinders by encapsulating bovine serum albumin (BSA) inside the inner PEGDA layer. By comparing the release characteristics with a single, unfolded PEGDA patch, they found that the folded structure released BSA at a much lower and steady rate, with an 80% lower BSA burst in the first 24 hours compared to the PEGDA patch. The folded structure retained up to 70% of the BSA even after 6 days, compared to the PEGDA patch which released most of the BSA content within the same period. Baek et al. thus demonstrated the effect of programmable and adaptive materials in presenting more barriers to the release of biomolecules, thus creating more steady release profiles.

Figure 3.5. Depiction of the hydrogel PEGDA layers with different crosslinking densities in the multi-walled tube. The release of encapsulated growth factors in uniaxial direction, which enhances vascularization, is depicted as well. Reprinted from Wiley with permission.85 4. Patterned Thin Films While the usage of homogenous polymeric thin film bilayers has shown promise in encapsulation of biomolecules, they are limited in the shapes that they can produce. These 23 ACS Paragon Plus Environment

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homogenous layers are only capable of rounded shapes like cylinders and capsules as it is considered practically impossible to create sharp angled 3D structures. The limitation in the shapes produced by the homogenous films is a disadvantage in designing drug delivery systems as the surface properties of the drug delivery packages are critical in determining their interactions with cells and proteins in the human body. To achieve sharp angled shapes, one method would be to create patterned thin films. A few possible approaches exist for making patterned thin films. One way would be to insert small portions of active polymers into the passive polymer to create hinges or folds as seen in Figure 4.1 below.71 As shown in Figure 4.1 in a method that is analogous to origami, the patterned thin films are capable of forming complex polyhedral-like 3-sided pyramids, 4-sided pyramids, cubes and hexagonal prisms. Even more complex shapes are possible by increasing the number of faces and hinges. Another approach to creating these containers is manipulating the shapes of bilayers in order to create templates that will also be able to fold and form capsule-like containers. These are typically flower shaped or rounded and will fold together to form spherical shape.

Figure 4.1. Folding of a patterned polymeric thin film (a) matellic 2D precursors, (b) solder covered 2D precursors and (c) folded polyhedral. Reprinted from the American Chemical Society with permission.71

4.1 Techniques for the Patterning of Thin Films 24 ACS Paragon Plus Environment


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Many of the techniques utilized for the patterning of thin films have been adopted from silicon based micropatterning. One such technique is photolithography, which has been adopted to create polymer patterns at both micro and nanoscale86. Photolithography works on the principle of exposing the material to electromagnetic radiation (UV light etc.) to change the solubility of the material. Photomasks can be employed to selectively expose only certain portions of the material to the electromagnetic radiation to selectively dissolve or maintain those portions upon developing of the material with a solvent. Many different polymers have been successfully patterned using photolithography,

including

biocompatible

polymers

like

PVA,

PMMA

and

SU-887.

Photolithography is currently one of the most developed and mature technologies used in the creating of patterned polymer thin films for programmable adaptive applications. One major disadvantage, however, is that it limits the polymers that can be used to photosensitive polymers. Electron beam lithography can also be used to pattern electron sensitive polymers like PMMA88. Upon impact with the surface of the polymer, free radicals and radical cations are generated in the surface layer. The solubility of these regions is thus changed compared to the bulk polymer. Application of a solvent will then selectively dissolve these regions and create patterns in the polymer thin film. Features as small as 50 nm have been achieved by electron beam lithography. However, it also has shortcomings in that it has very long writing times, making it impractical for scaling up to mass manufacture. In a similar vein, Ion beam lithography utilises ions instead of electrons to pattern the polymer films. A focused ion beam of ions like Ga+, H+, or He+ are fired at the polymer thin film and are able to penetrate with well-defined paths, with the penetration depth being dependant on ion energy. Similarly to Electron beam lithography, the ions trigger chemical reactions that affect the solubility of the polymer. Ion beam lithography has been shown to be capable of creating accurate patterns with well-defined walls, with high-aspect-ratio walls of 30 nm width with sub-3 nm edge smoothness89. However, the technique shares the same weaknesses as electron beam lithography in its long writing times and impracticality for scaling to mass manufacture. Nanoimprint lithography (NIL) is another technique that can create patterned polymer 25 ACS Paragon Plus Environment

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thin films at lower cost compared to more conventional lithographic techniques. In this technique, a thermoplastic polymer film is deformed by pressing a hard mold that contains the desired patterns onto it at controlled temperatures and pressures. The higher temperatures and pressure allows the thermoplastic polymer to flow into the mold. Subsequently, reducing the temperature and pressure reduces the viscosity of the polymer film and freezes the desired patterns in place, creating the patterned thin polymer film. In order for effective imprinting of the patterns, temperatures of 70 - 80 °C above the Tg of the polymer are required90. A wide variety of patterned polymers films have been successfully prepared using NIL with features as small as 5 nm. Some of these polymers include PMMA91, block co-polymers92 and conducting polymers93. A study by Guan et al.41 also demonstrated the use of a PDMS stamp to pattern PEGDMA, PVA and PLGA thin films. NIL has also been shown to create patterned features in polymers at the size scales as conventional photolithographic techniques. Arrays of imprints of 10 nm diameter patterns of PMMA on gold substrate or 6 nm diameter patterns of PMMA on silicon substrate have been achieved91. The results of the imprinting depend strongly on the Tg of the polymers, film thickness, width and height of imprinted features and the mechanical properties of the polymer94.While NIL has shown much promise in creating nanoscale patterns for polymer thin films, major problems remain. While most thermoplastics can, in theory, be patterned by the NIL process, most commercially developed thermoplastics are not optimized for the NIL process as most commercial thermoplastics do not readily release the mold upon imprinting. The polymers have a tendency to adhere to the mold, leading to defects in the patterning of the film95. There have also been challenges in scaling up the NIL technique due to the lifetime of the mold. Due to high temperature and pressure cycles, the average lifetime of a NIL mold is 50 cycles. In order to address this problem, room temperature NIL has been attempted on polymers like PMMA. However, the sizes of the imprints are of a larger order compared to conventional thermal NIL, with features on the order of 100 nm96.

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4.2.1 Patterned hydrogel films Shim et al.97 reported on the design of patterned hydrogel based bilayer that could fold and form spherical shapes (Figure 4.2). The polymers utilized were P(HEMA-co-AA) and P(HEMA) PHEMA, which were chosen in part due to their biocompatibility. P(HEMA-co-AA) acted as the active layer as a hydrogel as it could swell up to 1000% under basic conditions. A photolithographic technique was used to pattern the bilayer. Two different shapes of micro-particle were produced: a snowman-shaped and a flower shaped particle with 400 µm diameter. The folding of the microparticles was triggered by pH, with folding of the bilayer at higher pH and unfolding under acidic conditions. The folding process was found to take up to one minute at pH 9 while the unfolding process was found to take up to 10 seconds at pH 4. This process was found to be highly repeatable, with little damage to the polymer walls.

Figure 4.2. (a) Folding process of the two halves of patterned bilayer and (b) Photolithographic patterning process for bilayer. Reprinted from Wiley with permission.97 The encapsulation and release capability of the system was investigated through the capture and release of polystyrene particles. Releasing the containers into a suspension of 1 µm polystyrene at pH 9 was shown to be effective at encapsulating the polystyrene particles and holding them without any premature release. The permeability of the capsule was also investigated by testing of the diffusion of K-dextran particles. The capsules were found to be highly impermeable at pH 9, with very low diffusion rates of K-dextran particles. Releasing of these capsules into a solution of pH 4 was shown to cause the capsules to release their payloads. The possibility of site specificity through the inclusion of magnetic particles was also investigated. Shim et al. added iron (III) oxide particles into the passive layer during polymerization. The resultant capsules were found to be responsive to 27 ACS Paragon Plus Environment

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magnetic fields, aligning to an applied external magnetic field. This raises the possibility of a more site specific delivery of the capsules within the human body by applying a magnetic field to the diseased site of a patient. Fusco et al.98 also reported on a near infrared (NIR) responsive hydrogel bilayer system that was responsive to magnetic fields. An active 10-µm poly(ethylene glycol) diacrylate (PEGDA) layer was first polymerized and subsequently coupled with a poly(Nisopropylacrylamide) (PNIPAAM) based graphene oxide nanocomposite. Similar to Shim et al., photolithographic methods were used to pattern the bilayer. The templates formed were of crossshaped, jellyfish-like, and Venus flytrap-like structures. The bilayers would form 3D structure upon release, with differential swelling causing the folding of the bilayer. The 3D container was used to encapsulate an alginate microparticle with iron (III) oxide nanoparticles immobilized within. The alginate microparticles were intended to serve as porous drug reservoirs and the iron (III) oxide nanoparticles were included to allow control of the entire capsule using magnetic fields. Release of the payload was achieved by exposure to NIR, which caused the graphene oxide/NIPAAM layer to heat up and swell, causing actuation of the 3D structure. The capsule was also found to be stable up to 40 °C, meaning that they would not instantly release their payload at the temperatures present within a human body, allowing the possibility of delayed and controlled release of the payload. The possibility of magnetic manipulation of the capsules was also demonstrated using OctoMag, a targeted electromagnetic field generator designed for in vivo applications. The capsules were found to be manoeuvrable in figures of eight by the device and magnetic agitation was shown to help facilitate the release of the alginate microparticles upon opening of the bilayer capsule. Fusco et al. demonstrated the potential of this system for delivering biomolecules through an in vitro test of the encapsulation of D1 mouse mesenchymal stem cells (mMSCs) inside the magnetic alginate microbeads. Up to 80% of the cells were shown to be viable after seven days. A second experiment with encapsulation of brilliant green, a dye that also acts as a topical antiseptic, showed that the bilayer was effective at presenting a physical barrier to biomolecules encapsulated within the alginate. The effect of the duration of NIR exposure on the rates of diffusion of brilliant green was 28 ACS Paragon Plus Environment


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also investigated, with short exposures found to lead to lower diffusion of brilliant green. In all, Fusco et al. demonstrated the possibility of using their bilayer system to encapsulate a biomolecule or cell, direct it to the diseased site of a patient using magnetic fields and trigger the release of the drug using NIR. 4.2.2 Polymeric Programmable Adaptive Microactuators Microactuators for drug delivery are a field in which programmable adaptive polymers have been utilised to a great extent. Through the use of lithography and other shaping techniques, micro sized polymer bilayer films are patterned into claw like shapes that are designed to improve targeted drug delivery by allowing the polymeric arms to fold and dig into the mucus linings and uneven cell walls, improving adhesion to specific disease sites. This allows the drug delivery device to release its payload of drugs to the site in a much more specific manner. A second approach would be to make the microactuators responsive to magnetic fields to allow them to be directed to specific parts of the body. They would be used to hold polymeric microbeads that act as drug reservoirs. Upon reaching the targeted site, some form of trigger, like pH change or infrared radiation can be applied to trigger the release of the drug reservoir by making the microactuator unfold. pH trigger has been utilised for many different drug delivery applications.99-104 He et al.81 have also demonstrated a pH responsive programmable adaptive patterned hydrogel system that was designed for targeted drug delivery in the form of a mucoadhesive drug delivery systems (MDDS). MDDS are drug delivery systems that are designed to adhere to a diseased site in the mucus lined intestines and steadily release drugs in a directed fashion to the site. They utilized a finger-like bilayer that consisted of a PMAA based pH sensitive hydrogel and a non-swelling layer of PHEMA. A third mucoadhesive layer was attached on top of the bilayer consisting of the model drug, A08, mixed with Carbopol 934 and PVA. The folding properties of the bilayer was shown to improve mucoadhesiveness as the bilayer was able to “curl” into the mucus. The improved performance was demonstrated in vitro through a flow test, where the folded devices were seen to remain attached to a intestinal-like mucus surface for more than 90 minutes compared to 60 minutes for a PCL patch and 10 minutes 29 ACS Paragon Plus Environment

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for a PHEMA patch without folding capabilities. The force required to detach the samples from the mucosal surface was also investigated, with the folded device having the highest detachment force compared to PCL, PHEMA and Carbopol 934/PVA patches. The release profile of the model drug was also investigated. The folded device was shown to have a more steady release profile for the model drug compared to PCL, PMAA patches. He et al. hypothesized that the PHEMA layer in the folded device provided a greater barrier to A08 initially, leading to more sustained release. The delivery of the model drug through the mucosal layer was also investigated. He et al. demonstrated that the folded device managed to deliver a higher percentage of A08 across the mucosal layer compared to the PCL patch and a pure solution. A similar test was done for BSA as a model drug which led to similar results. He et al. thus managed to demonstrate the potential of the programmable adaptive PMAA/PHEMA bilayer for drug delivery to intestinal, ocular, vaginal, buccal and vaginal sites due to its improved mucus adhesion and improved drug transport and delivery capability compared to non-folding alternatives. Li et al. 105 demonstrated a pH responsive microactuator for drug delivery that could be magnetically controlled using a bilayer of PHEMA as the swelling hydrogel layer and PEGDA. Iron (II, III) oxide particles (Fe3O4) were embedded inside the PEGDA layer to make the device controllable with magnetic fields. The layers were fabricated using a conventional photolithographic technique. The device was used to grip a PCL microbead containing the anti-cancer drug docetaxel (DTX). The device was shown to fully close at pH of 9.58 and open at pH of 2.6. The ability of a magnetic field to manipulate the device was demonstrated by moving the device using the generated magnetic field of an Electromagnetic Actuator system with a maximum moving velocity of 600 µm s−1. The tumour killing ability of the system was also investigated through in vitro tests on mammary carcinoma cells (4T1). This system was shown to decrease 4T1 viability by 29%.

Breger et al.

106

demonstrated a thermal and

magnetically responsive patterned hydrogel system to create a “gripper” that is claw shaped shape. Patterned bilayers consisting of a stiff PPF layer and a soft P(NIPAM-AAc) hydrogel layer were used. Similar to the previously discussed studies, iron (III) oxide nanoparticles are included in the 30 ACS Paragon Plus Environment


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hydrogel layer to make the container responsive to magnetic fields. A photolithographic method was used to pattern the layers into a star shaped template as shown in Figure 4.3 below.

Figure 4.3. Reversible programmability and adaptivity of a soft “gripper”. The PPF segments are coloured dark blue while the P(NIPAM-AAc) layer is coloured light blue. Reprinted from the American Chemical Society with permission.106

The opening and closing of the grip is triggered by temperature. Above 36 °C, the P(NIPAMAAc) layer ejects water and contracts, causing the grippers to first open and then close such that the PPF segments face outward. Below 36 °C, the P(NIPAM-AAc) absorbs water and swells which causes the microgripper to open and then close in the opposite direction such that the P(NIPAMAAc) layer faces outward. The optimal for the thickness of the hydrogel layer was found to around 45 µm for a PPF layer of 10 µm. The magnetic responsiveness and gripping capability of the system was demonstrated by the encapsulation of fibroblast cells. The bilayer was directed towards a clump of fibroblast cells using directed magnetic fields and was observed to grip and excise cells from the clump. The gripper on the cells was found to be stable at 4 °C and the cells were observed to be viable. Upon raising of temperature to 37 °C, the gripper released the fibroblast cells. This thus lead Breger et al. to propose this bilayer grip system as a drug delivery mechanism for live cells, with cells stored at low temperature. Directed injection of these hydrogel bilayer grippers with the cells into the human body would cause the release of the cells. Magnetic fields could also be employed to target specific diseased parts of the human body. The use of these grippers for drug delivery was further studied by Malachowski et al.107 in both in vitro and in vivo experiments. Using a similar bilayer design to Breger et al, Malachowski et al. examined the capture and release of E. coli strains 31 ACS Paragon Plus Environment

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TG1, TG2 and TG3 by the bilayer grip gripper. The grippers were made using a similar photolithographic approach and soaked in solutions of the bacteria strains. Their release profile was then studied. They found that the release of TG1 and TG2 was much faster than that of TG3, with most of TG1 and TG2 being released within 6 hours while TG3 was released over a period of 7 days. This led Malachowski et al. to hypothesized that a mixture of these two release profiles could be utilized for biphasic drug release. The encapsulation and release of the anti-inflammatory drug mesalamine and the chemotherapeutic drug doxorubicin (DOX) at 37 °C was also examined. A complex of mesalamine-TG3 was released steadily by the gripper over a period of 15 days while a complex of DOX-TG1 had a faster release profile of about 1 day. The role of the gripper in increasing site specificity was also investigated. The gripping action was thought to be able to allow the bilayer to grip onto cells in the human body and allow more effective delivery of drugs. This effect was tested by comparing a DOX infused gripping bilayer with a non-gripping DOX infused patch of similar size. The gripping bilayer was found to kill MDA-MB-231 breast cancer cells more effectively than the non-gripped patch in vitro at 37 °C. In vivo tests were also carried out by Malachowski et al. TG1 grippers loaded with blue dye were inserted into the stomach of a pig through an endoscope. The grippers were observed to grip onto the stomach walls of the pig upon delivery and the model drug was visually observed to be delivered into the walls of the pig’s stomach through the blue dye. 4.2.3 Hinged programmable adaptive polymeric containers Azam et al.

108

reported on the design of a programmable adaptive patterned all polymer system

consisting of SU-8 faces and PCL hinges (Figure 4.4). 2D templates of the SU-8 faces and hinges were prepared with a combination of photolithographic and lift-off deposition on top of a PVA sacrificial layer. The PCL regions act as hinges and cause the template to spontaneously assemble into a 3D structure at 58°C. This is driven by the liquefying of PCL and a change in shape to minimize surface tension of the PCL after its phase change. To encourage precise angles of 90°, locking hinges were designed to overlap and fuse to prevent panels from overfolding. Upon cooling, 32 ACS Paragon Plus Environment


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the structures were locked in place as the PCL solidified. Azam et al. reported an 18% success rate at creating well-formed cubes.

Figure 4.4. Process of creating patterned SU-8/PCL programmable adaptive material-based container. (ai) SU-8 panels are patterned using conventional lithography. (aii) PCL is deposited in gaps to form hinges. (aiii) Sacrificial PVA layer is dissolved in water. (b i-iii) Schematic showing the folding of the template. (c i-iii) Actual folding process of 100 micrometer length cube. Reprinted from Springer with permission.108 Azam et al. also investigated the potential of these 3D structures as a method of encapsulating and delivering biomolecules through the encapsulation of various biomolecules like fibroblasts cells, mammalian pancreatic cells and bacteria. Loading of these cells into the structures was tested using two methods. Firstly, bacteria was loaded into the microstructures by the tumbling of the structures with small gaps in the hinges through highly concentrated bacteria containing solutions. Secondly, mammalian fibroblast cells were coated with fibronectin and attached to the 2D templates before folding. Upon heating to 58 °C, the templates folded into 3D cubes, encapsulating the cells. The viability of the bacteria and cells was studied. The bacteria was observed to go through several growth cycles over one day. The cells were also observed to form clumps and continue multiplying over the course of seven days. The release characteristics of the containers were also studied. The PCL hinges were observed to degrade over the course of 30 days. The pH of the environments were was shown to affect the degradation profile of PCL, with higher pH leading to greater degradation. Azam et al. proposed the use of co-polymerization techniques for the PCL in order to effectively control the degradation profile of the PCL. Release of contents of the containers was observed both 33 ACS Paragon Plus Environment

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through the pores in the container walls and the opening of the container upon degradation of the PCL hinges. This thus highlights the potential of the Su-8/PCL microstructure in controlled drug delivery. Similarly, Yoon et al.109 reported on the design of a programmable adaptive hinged system making use of P(NIPAM-AAc) (Figure 4.5). The P(NIPAM-AAc) films were patterned using a photolithographic technique utilising UV light. The panel sections were exposed to strong UV light to encourage extensive cross-linking. The film was then exposed to a solvent that dissolved the non-crosslinked portion. Additional P(NIPAM-AAc) was inserted between the panels to act as hinges and exposed to low energy UV light to cause it to weakly cross linked. The hinges are therefore softer and less stiff than the panel segments. The templates will then fold and form 3D structures upon heating and at low pH. Yoon et al. successfully created pyramid and cubic containers of around 500 µm and demonstrated the encapsulating capabilities of the containers by encapsulating small beads. They also created gripper systems with six arms and three joints which close upon raising the temperature to 65 °C and open upon lowering of temperature to 25 °C.

Figure 4.5. Steps involved in creating P(NIPAM-AAc) films with hinges (A) Panel: high energy UV exposure, (B) highly crosslinked panels, (C) hinger: low energy UV exposure, (D) crosslinked gradient along the hinge thickness and (E) pH and thermal stimuli. Reprinted from Springer with permission.83

4.2.4 Patterning of thin film with UV light



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Jamal110 reported on a method to induce programmable and adaptive functionalities in a single polymeric SU-8 layer by selectively exposing the SU-8 film to UV light. As SU-8 is a negative photoresist, the extent of crosslinking in the film can be controlled by selectively altering the magnitude of the UV light exposure energy. This process of photo patterning of the film can create a cross link gradient throughout the film and causes a stress gradient to form upon exposure of the hydrophobic SU-8 film to water solvent. The less cross-linked portions of the film contracts to a smaller extent then the more cross-linked portions and this leads to folding. The films were thus observed to fold into cylindrical shapes with radiuses in the order of micrometres upon solvating in water. The radius of curvature was found to be dependent of the exposure energy of the UV radiation, with higher exposure energy leading to larger radiuses. The formation of more complex structures was also achieved through the use of photomasks to selectively create high and low cross-linked regions. Hinged shapes and cubes with flat faces were achieved through this process.

5. Shape Memory Polymer Systems Shape memory polymer systems are unlike bilayered or patterned programmable adaptive systems previously described in many ways. While these polymer systems also are stimuli triggered, the change in the shape of these systems are is due to the ‘lock’ in its cross-linking, molecular entanglement or interpenetrated network. This lock allows the transition from its temporary shape to its original or permanent shape which can be attributed to the polymer’s crystallization or melting transition, vitrification or glass transition, anisotropic or isotropic transition, reversible molecular crosslinking or supramolecular association/dissociation. These are demonstrated in Figure 5.1.


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Figure 5.1. Different molecular structures and transitions of shape memory polymers are depicted. Reprinted from Elsevier with permission.111 The advantages of shape memory polymer systems are vast. These systems are typically lightweight, allowing for far higher elongation of more than 200%, high shape recovery properties, possess a range of glass transition temperatures, are both biocompatible and biodegradable, have low cost and are easy to process. Although shape memory polymers can be used in a wide variety of biomedical applications such as sutures which can self-tighten, self-expanding stents, neuronal electrodes and thrombectomy devices111, here we are mostly interested about drug encapsulation and delivery systems, most commonly, drug releasing stents and degradable implants which also allows the elute of drugs. Both physical crosslinking and chemical crosslinking exist for shape memory polymer systems. Physically crosslinked copolymers are frequently made up by linear polymer blocks. These blocks can either be hard, which helps the system to attain its permanent shape, or be switching segments, which allows its temporary shape to form. Intelligent drug delivery systems have to undergo ‘programming’, a term to describe the fixation of its temporary shape. Upon implantation and being exposed to physiological environments, it then can recover its original shape. This is shown conceptually in Figure 5.2.



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Figure 5.2. The process of programming a shape memory polymer and its eventual shape recovery and drug release for biomedical applications. Reprinted from Elsevier with permission.112 The most widely studied shape memory polymer systems utilize thermally triggered transitions where the locks in the network can be covalently-bonded joints or physical in nature such as crystallites112. Another characteristic necessary is the presence of switching segments in the polymer chain in forming amorphous domains. To attain the permanent shape of the shape memory polymer, the polymer is heated above its melting temperature Tm or glass transition temperature Tg to allow the long range chain segmentation motion to occur, giving the polymer flexibility. When the polymer is cooled below Tm or Tg,, the shape would then be fixed to that of its temporary state through the solidification of its switching domains by crystallization or vitrification. If the shape memory polymer is then heated above a temperature known as its switching temperature Tsw, it can then recover its original shape. This process is fundamentally governed by the gain of entropy. Although systems based on crystallisable switching segments are more common, the advantages of amorphous shape memory networks have its advantages. For instance, if the entirety of the polymer system is amorphous, it can achieve more homogeneous degradation and provide a more consistent load support and drug release. Also, as there is no need to preserve crystallinity, it allows for greater variations of copolymer for the system. Lastly, the more densely packed crystalline segments in the system can hinder the release of drugs and as such, a fully amorphous system can achieve higher drug loading. In general, there are two main methods to encapsulate the drugs into the network system. The first involved the drugs to be loaded in the network precursors before being crosslinked and the second, to deposit the drugs after crosslinking where the network has already formed. This 37 ACS Paragon Plus Environment

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is often done concurrently with swelling by placing it in a drug solution and drying it after encapsulation. Also, the drug particles can be either dispersed throughout the entire polymer implant or only specific regions, most commonly, its surface. Due to the shape memory ability of the polymer system, drug distribution is important as upon implant, a burst release of drugs is highly plausible113 which may have negative effects on the surrounding human tissue. 5.1 Temperature Sensitive Shape Memory Systems Temperature responsive polymeric shape memory systems have great potential in minimally invasive surgery.114,

115

These systems can be placed into its desired location when it is in its

temporary compressed shape through a small incision. Upon heating up to the body temperature, past Tsw, it could attain its permanent shape, all the while releasing drugs into the implant site. These types of shape memory polymer systems require a highly narrow transition temperature range slightly above the body temperature however. Nagahama et al.

116

developed and tested the shape

memory properties of star shape branched oligo(ε-caprolactone) as the crystalline phase and hexamethylenediisocyanate as the crosslinker (Figure 5.3). Poly(ε-caprolactone) (PCL) has a melting temperature of 60 °C, is biodegradable and acts as the thermally reversible portion in the structure. Nagahama et al. utilized relatively short PCL chains to form the branched system in order to lower Tm and the resulting system has a very narrow shape recovery temperature range only across 2 °C from 37 to 39 °C which was noted to be advantages as it could prevent the implant from shifting from its implant site and it could decrease the trauma to cells during implantation. In the study, the drug loaded, theophylline, was achieved during the crosslinking process and the system managed to sustain the release of the loaded drug for over a month without any noticeable initial burst release in a phosphate buffer solution. Even with the loaded drug, Nagahama et al. still observe the PCL shape memory system to be soft and flexible enough for shaping.



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Figure 5.3. (a) shows the permanent spiral-like shape of the poly(ε-caprolactone) shape memory system. (b) shows the temporary rod-like shape fixed at 25 °C and (c) shows the recovered shape after being in contact with water at 42 °C. Reprinted from the American Chemical Society with permission.116 In another example closer to in-vivo implementation, Wache et al.117 demonstrated a polymeric shape memory stent. Although shape memory alloy stents based on nickel and titanium or nitinol are already in the market today, benefits polymeric stents have are its biocompatibility and possible biodegradability, and its drug encapsulating and eluting properties. The shape memory stent was built from thermoplastic polyurethane which consist of a frozen crystalline phase and a reversible amorphous phase. The transition is thus dependent on the temperature being between the Tm and the Tg for its permanent shape and below Tg for its temporary shape. For Wache et al., the loading of drugs was made after polymerization and during processing through injection molding and dip coating. Results for the drug release showed that a continuous release at low drug concentration was possible. Wishke et al. in two separate papers112,

113

, published details on a degradable shape

memory system for controlled drug release based on a copolyester urethane network and oligo[(εcaprolactone)-co-glycolide] dimethacrylate. For the copolyester urethane system, its synthesis was achieved

through

ring

opening

polymerization

of

star-shaped

oligo[(rac-lactide)-co-

glycolide]tetrole. This was then crosslinked with 2,2,4- and2,4,4-trimethyl-1,6-hexamethylene diisocyanate to form an amorphous polymer network with a high gel content. The programming of the shape memory system was achieved in both dry and aqueous environments although it was found that the switching temperature TNO was reduced in aqueous environments. When Wishke et al. tested the system after the loading of drugs, which was incorporated into the system during the swelling process, it neither affected the mechanical properties compared to the drug-free system,


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nor affect other shape memory properties except for the swelling ratio which increased. This may be a result of osmotic effects. The burst release of drugs was observed to be not more than 20% when using enoxacin or ethacridine lactate, and was even less when the system was loaded with nitrofurantoin. Past this burst release, ethacridine lactate and nitrofurantoin displayed a linear and highly controllable release while enoxacin remained mostly trapped within the shape memory polymer system. Past 30 days, the system began to degrade causing a large increase in the rate of release of all three drugs. Interestingly, the release of ethacridine lactate from its programmed temporary shape showed only a slight difference from its non-programmed permanent shape and this may be due to the reduction of the free volume of the polymer system. The more linear conformation of the polymer chains would then hinder the release of drug out of the polymer matrix. Wishke et al. also analyzed the drug distribution after loading in the copolyester urethane system. Drug crystallization where the drugs aggregate to form crystallites was observed for all three drugs loaded which may provide a reason why their observed burst release is lower than others which have a higher drug concentration on the surface. This also provided an explanation to their observations of the copolyester urethane system having no change in Tg after the loading of the drug, where a lower Tg was expected due to the higher free volume from the presence of small molecules. They also pointed out that the drug loaded into their amorphous system was far higher than that of partially crystalline systems by more than two times. 5.2 Other Stimuli Responsive Shape Memory Systems Although temperature is the most common triggers for shape memory systems, there do exist others stimuli which can also induce the shape transition of these systems. Here, pH and solvent triggers will be discussed. Han et al.118 proposed and synthesized a pH reversible crosslinked β-cyclodextrin modified alginate and diethylenetriamine modified alginate which keeps its temporary shape at pH 11.5 and recovers its permanent shape at pH 7. Similar to the two phases in thermally induced shape memory systems, the fixed phase in this pH controlled system consisted of crosslinked alginate chains and calcium cations which are stable at neutral pH. The temporary switching phase was 40 ACS Paragon Plus Environment


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made from the crosslinked β-cyclodextrin and diethylenetriamine, which can dissociate at different pH levels. It was found that the formation of inclusion complexes between the β-cyclodextrin and the short chain diethylenetriamine caused the pH responsiveness of the system. In a basic medium, the alginate chains and the crosslinked β-cyclodextrin and diethylenetriamine are fixed. As the pH decreases, the protonation of amino groups in diethylenetriamine induces the inclusion complexes between β-cyclodextrin and diethylenetriamine dissociates, rendering the system to be soft and easily deformed. The inclusion complexes thus serve as the switches in this system (Figure 5.4).

Figure 5.4. Mechanism of the shape memory effect shown for the crosslinked β-cyclodextrin (βCD) modified alginate (Alg) and diethylenetriamine (DETA) modified alginate system. Its temporary shape can be seen at pH 11.5 and its permanent shape at pH 7.0 due to the formation of inclusion complexes (ICs). Reprinted from Wiley with permission.118

In another pH based study which utilizes acidic triggers rather than basic as per the previous example, Chen et al.119 synthesized a structure consisting of pyridine rings and polyurethane and tested its drug delivery capabilities. The mechanism for the pH triggered transition is due to the hydrogen bond between the N atom in the pyridine ring and the H-N of the urethane, which is present in neutral or high pH but is disrupted at low pH. This acts as the switch for this system for the transition between its temporary and permanent shapes. For pyridine, a Lewis base, it readily accepts H+ and becomes protonated in acidic environments. This leads to an increase in hydrophilia and water molecules could then diffuse into the polymer, causing it to swell. However, in basic 41 ACS Paragon Plus Environment

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environments, the N atom will be deprotonated and the polymer would lose water molecules instead. Unlike the previous examples of drug delivery systems based on thermally activated shape memory polymers, Chen et al. encapsulated drugs more similar to the bilayer folding method. In physiological environments, the release of drug was seen to be extremely low and rapidly increased in acid environments. It should be noted that this method is different from other pH systems that irreversibly breaks down due to the cleavage of pH-sensitive bonds in the system. A solvent based shape memory system was demonstrated by Xiao et al.

120

where sulfonated tetrafluoroethylene

based fluoropolymer-copolymer, or Nafion, was shown to display shape recover at different solvents analogous to the temperature or pH memory effect. The solvents tested were water, acetone, ethanol and isopropyl alcohol. Programming of the Nafion-based system into its temporary shape was made in air, at different temperatures. The shape recovery of Nafion under different solvents is shown in Figure 5.5.

Figure 5.5. The graph shows the effects of solvent on the shape memory properties of Nafion. The strain recovery is shown for specimen programmed at 160 °C and 140 °C across different solvents. Reprinted from the Royal Society of Chemistry with permission.120

The solvent-based memory effect is due to the tetrafluoroethylene backbone of Nafion having perfluorovinyl ether with a terminal sufonic acid group. Two different swelling envelopes exist for the backbone and the ionic group where their degree of swelling was shown to be different. Also, solvents behave differently upon contact with Nafion. Alcohol solvents preferentially migrate into the non-ionic portions of the polymer, i.e. other portions besides the ionic sulfonic acid group. This occurs due to the stronger van der Waals interactions between the methyl group of the alcohol and 42 ACS Paragon Plus Environment


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the crystalline tetrafluoroethylene backbone. It should be clarified that the shape memory transition is still inherently linked to the glass transition temperature Tg, and the role of the solvent is to increase the chain mobility, which lowers the Tg to make it appear as though the solvent is the trigger. When the solvent is absorbed by the polymer chain, the configurational entropy resulting from the mixing of polymer and solvent molecules increases. The solvent with the highest swelling ratio would cause a greater shape recovery due to the larger decrease in viscosity and relaxation time, driving the configurational entropy up. A simpler but possibly more effective solvent-based system was developed by Nöchel et al.121 who utilized a hydrogel containing hydrophobic telechelic PCL diisocyanoethyl methacrylate as the crosslinker and hydrophilic poly(ethylene glycol) monomethyl ether monomethacrylate as a co-monomer, which displays shape memory transition upon contact with aqueous environments (Figure 5.6). Similar to the above example, the transition or switching temperature TSW is the trigger for the swelling and subsequent shape transition. However, TSW can be influenced by water molecules, which act as plasticizers and lower the Tg of the switching domains and promotes softening. Nöchel et al. found that by lowering the PEG ratio in the system, the degree of swelling lowers accordingly. In general, the increase in either of the polymer species would lead to an increase in the crystallinity of that polymer segments and the properties of the shape memory system would change accordingly. At higher temperatures, the PCL has greater mobility, which leads to the decrease in water uptake due to the hydrophobic nature of this chain. Also, when the PEG portions swell upon water uptake, i.e. in an aqueous environment, mechanical stress is applied to the PCL phase which results in its lower crystallinity, where smaller crystallites lead to smaller Tm . Regarding crystallinity, Nöchel et al. found that the as the crystallinity of PCL increases, the orientation of the chains develop into lamellae-like structure, which has orientation in the longitudinal direction. This behaviour thus affects the switching ratio of the hydrogel system. However, upon decreasing the PCL ratio below a certain threshold, insufficient numbers of physical crosslinking led to the loss of the shape memory property.


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Although the drug loading properties of this system was not extensively studied, Nöchel et al. proposed this system as a potential controllable drug release system.

Figure 5.6. A depiction of the hydrogel system containing hydrophobic telechelic poly(εcaprolactone) diisocyanoethyl methacrylate (PCLDIMA) and hydrophilic poly(ethylene glycol) monomethyl ether monomethacrylate (PEGMA). The amorphous and crystalline PCL chains are shown in red and blue respectively. The amorphous PEG chains are shown in grey. The diagram shows the swelling, deformation and programming of the system. Reprinted from Elsevier with permission.121 5.3 Potential Applications for Shape Memory Systems 5.3.1 Intravascular Stent As previously discussed, the study made by Wache et al. proposed a polyurethane shape memory system for drug delivery, which utilizes the different properties of its amorphous and crystalline phases. Although other studies on shape memory polymeric intravascular stent did not highlight their drug encapsulation and delivery capabilities, they do have potential for this application as well. Tert-butyl acrylate, and poly(ethylene glycol) dimethylacrylate has been proposed by Gall et al.122 as possible material for stent applications where they determined that 10 wt.% crosslinking allows the programmed shape to exist at room temperature. This system will revert to its permanent shape upon being heated to body temperature. Chen et al.123 made a study on chitosan/PEO/glycerol for the synthesis of biodegradable stents. The trigger for this system is solvent-based and will activate upon contact with water. These same authors also proposed a genipin and chitosan crosslinked system which has similar functionalities and was tested with a loaded drug sirolimus. This system 44 ACS Paragon Plus Environment


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was found to be mechanically stronger compared to the chitosan/PEO/glycerol-based system. Another possible biodegradable shape memory stent was studied by Venkatraman et al.

124

using

PLLA and PGA, which can expand at body temperature. 5.3.2 Aneurysm Treatment Aneurysm refers to a localized swelling of the artery wall due to the weakening of the wall and can exist in any blood vessels, although most commonly found in aortic, cerebral, popliteal, mensenteric, and splenic blood vessels. Shape memory polymers can help to treat these conditions in a method similar to using stents for atherosclerosis, and can also incorporate drug delivery methods as well. These polymers are inserted and expand outwards, giving them the role of filling devices in the artery. Wilson et al.125 demonstrated a hexamethylenediisocyanate, N,N,N’,Ntetrakis(2-hydroxypropyl)ethylenediamine, and triethanolamine, which was shown to be able to treat a 10-mm aneurysm through in vitro testing. Wong et al. 126 showed that Calo-MER™, a shape memory thermoplastic developed by The Polymer Technology Group, can spread inside a simulated aneurysm model without experiencing hemodynamic forces. Other shape memory foams include hexamethylenediisocyanate, triethanolamine and tetrakis(2-hydroxyl propyl)ethylenediamine developed by Maitland et al. and polyurethane developed by Maitland et al. and Raymonda et al.127. For the non-metagenic and poorly thrombogenic polyurethane-based system, it was found that the cytotoxicity is low and the addition of 4% tungsten improved radioopacity. 5.3.3 Tissue Engineering Xie et al.128 demonstrated a strong electroactive shape memory system based on polylactide and aniline trimer, with the polylactide forming ‘arms’ giving rise to a shape of a star. The system was found to have excellent shape memory properties such as, short recovery time, high recovery ratio and high fixing ratio, along with properties necessary for scaffolds such as high mechanical strength and biodegradability. Nardo et al.129 synthesized scaffolds based on solvent casting of gelatin microspheres and casting polyether-urethane on the compacted microspheres, which displayed shape memory effect without cytotoxic effects. 45 ACS Paragon Plus Environment

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6. Conclusion and Outlook This review summarized the methods for designing programmable adaptive materials including shape memory polymers, patterned thin film and differently stressed thin film as well as the advantageous applications of the programmable adaptive materials as carrier for drug delivery. Programmable adaptive materials belong to a class of stimuli-responsive materials known as smart polymers, which can change their shapes from the 2D flat/planar structures to 3D structure with various morphologies with the small or modest variations in their physical environment such as temperature, pH, magnetic field etc. This unique characteristic endows programmable adaptive materials with great promising applications in biomedical fields, especially as carriers for controlled drug release. However, there are also some concerns regarding the using of programmable adaptive materials in clinical applications. For instance, the miniaturization of programmable adaptive polymeric containers to nanoscale levels is still a challenge due to the inadequacy of current planar lithographic methods. Current lithography methods also do not allow the synthesis of many biocompatible and biodegradable polymers systems and as such, is limited with regards to selectivity. Also, the mechanical properties of programmable adaptive systems, a parameter vital for these systems, are still difficult to manipulate to a very precise degree. Manufacturability of these systems may also be a challenge. Due to patterning methods being highly parallel, large scale manufacturing and processing will most likely be expensive, far more than traditional colloidal methods. There may also be issues with the effects of the physiological environments with stimuli. All three methods discussed require a stimulus to achieve its drug release or mechanical support functionality. However, most studies tested these programmable adaptive systems in vitro and as such, physiological environments with many combination of stimuli may cause the folding to occur unpredictably. Much further study is needed in clinical tests in vivo before these proof-of-concept demonstrations can be adapted to real world usage. Another area which was not discussed in-depth in this review but has great potential is the field of tissue engineering. Biodegradable programmable adaptive scaffolds where cells can adhere 46 ACS Paragon Plus Environment


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For Table of Contents Use Only

A Review of Adaptive Programmable Materials and Their Bio-applications

Xiaoshan Fan1, Jing Yang Chung3, Yong Xiang Lim3, Zibiao Li2*, Xian Jun Loh2,3,4*

1

School of Chemistry and Chemical Engineering, Henan Normal University, China

2

Institute of Materials Research and Engineering (IMRE), A*STAR, 2 Fusionopolis Way, Innovis,

#08-03, 138634 Singapore 3

Department of Materials Science and Engineering, National University of Singapore, 9

Engineering Drive 1, Singapore 117576, Singapore 4

Singapore Eye Research Institute, 11 Third Hospital Avenue, Singapore 168751, Singapore

Table of Contents graphic (TOC)


Review of Adaptive Programmable Materials and Their

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