Design of Hydrogels for Biomedical Applications.

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Design of Hydrogels for Biomedical Applications Hiroyuki Kamata, Xiang Li, Ung-il Chung, and Takamasa Sakai* Flory first established the classical theories of polymers back in the 1940s,[2] which has helped the development of polymeric materials both from scientific and industrial points of view. Applications of polymers as functional materials were pioneered by Katchalsky et al.[3] In 1970s, the volume phase transition of polymer gels was well characterized by Tanaka et al., where polymer gels abruptly change their volume in response to external stimuli, such as solvent composition,[4] temperature,[5] and pH.[6] The discovery has triggered the advancement in the design of functional gel materials,[7] such as actuators, stimuli-responsive drug delivery systems, and sensors. While many research efforts had been devoted to the design of functional hydrogels in the 20th century, it has been widely believed by the public that hydrogels are mechanically brittle, and thus they easily fail to operate under a high mechanical stress. Since the early years of the 21st century, the mechanical properties of hydrogels have been significantly improved via novel designs of physical architectures of polymer networks; gels with movable cross-linkers,[8] energy dissipation mechanisms,[9] inorganic clays,[10] and homogeneous networks[11] are of successful examples. To focus on the preparation steps of hydrogels, there are two major methods: 1) polymerization of monomers with cross-linking molecules via free radical or controlled[12] polymerization, and 2) physical or chemical cross-linking of prepolymers via temperature-dependent sol–gel transition,[13] UV irradiation,[14] click chemistry,[15] and so forth. For biomedical applications, the latter may be particularly useful because the methods often provide hydrogels with injectability; such types of hydrogels are called injectable hydrogels,[16] which can be handled as aqueous solutions in the preparation stage, and considered key tools to allow for future non-invasive surgery. Despite their potential as biomaterials, the practical applications of hydrogels are often hindered by the difficulty in controlling their temporal change in shape after the installation (Figure 1). Typically, hydrogels in a water-rich environment such as in the body undergo the following three stages. First of all, hydrogels start to swell immediately after the installation, and continue to operate after they reach the equilibrium (Stage 1, Figure 1A); conventional hydrogels, which are made of hydrophilic components, essentially absorb water in an aqueous environment, increasing their volume (i.e., swelling). Afterward, the hydrogels that finished playing their role are required to degrade. In this stage, the degraded polymer networks lose

Hydrogels are considered key tools for the design of biomaterials, such as wound dressings, drug reservoirs, and temporary scaffolds for cells. Despite their potential, conventional hydrogels have limited applicability under wet physiological conditions because they suffer from the uncontrollable temporal change in shape: swelling takes place immediately after the installation. Swollen hydrogels easily fail under mechanical stress. The morphological change may cause not only the slippage from the installation site but also local nerve compression. The design of hydrogels that can retain their original shape and mechanical properties in an aqueous environment is, therefore, of great importance. On the one hand, the controlled degradation of used hydrogels has to be realized in some biomedical applications. This Progress Report provides a brief overview of the recent progress in the development of hydrogels for biomedical applications. Practical approaches to control the swelling properties of hydrogels are discussed. The designs of hydrogels with controlled degradation properties as well as the theoretical models to predict the degradation behavior are also introduced. Moreover, current challenges and limitation toward biomedical applications are discussed, and future directions are offered.

1. Introduction The design and development of biomaterials are critically important to our high standard of living. Since the mid-20th century, hydrogels have begun to play a significant role in the biomaterials field. Hydrogels are often defined as threedimensional polymer networks that contain a large amount of water, and possess two indispensable aspects as biomaterials:[1] 1) hydrogels are soft and elastic materials, of which the properties are shared by those of our soft tissues, and 2) hydrogels are open systems, which can exchange substances and/or energy between the inner and outer phases. Gel-containing products, such as soft contact lenses, wound dressings, and sanitary goods have already been introduced as daily commercial products.

H. Kamata, X. Li, Prof. U.-i. Chung, Dr. T. Sakai Department of Bioengineering School of Engineering University of Tokyo 7–3–1 Hongo, Bunkyo-ku, Tokyo 113–8656, Japan E-mail: [email protected] Prof. U.-i. Chung Center for Disease Biology and Integrative Medicine School of Medicine University of Tokyo 7–3–1 Hongo, Bunkyo-ku, Tokyo 113–0033, Japan

DOI: 10.1002/adhm.201500076

Adv. Healthcare Mater. 2015, DOI: 10.1002/adhm.201500076

© 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

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the elastic restorative force, allowing for a further increase in volume, which is called degradation-induced swelling (Stage 2, Figure 1B).[17] Continued swelling ultimately leads to the bulk degradation where the hydrogels become a sol (Stage 3, Figure 1C). Among the three stages, Stage 1—an unintentional increase in volume that happens before the programed degradation—is particularly troublesome in the medical context; the morphological change causes not only the slippage from the installation site but also local nerve compression.[18] Furthermore, swelling inevitably causes a critical decrease in the mechanical properties as expected in the classic polymer physics.[19] Thus, it is obvious that the swelling of hydrogels in an aqueous medium should be suppressed or preferably precisely controlled for a prolonged period. Another challenge will be to establish countermeasures for degradation-induced swelling (Figure 1B). In some biomedical applications, such as artificial locomotor system, hydrogels should play their role semi-permanently, competing against oxidative and/or mechanical stress. To take an example of scaffolds for regenerative medicine, hydrogels are required to degrade with a programed mechanism, synchronized with the tissue regeneration. This Progress Report provides a brief overview of the recent progress in the development of hydrogels for biomedical applications. The first topic includes the fundamental analysis of the dynamic swelling and shrinking properties as well as the introduction of practical approaches to design and fabricate injectable hydrogels with controlled swelling behavior. The second topic comprises the discussion about the design of the hydrogel with precisely controlled degradability, modeling of degradation, and experimental verification. The third topic focuses on the modeling of non-specific degradation, and experimental verification by accelerated degradation testing. Moreover, current challenges and limitation toward biomedical applications are discussed, and future directions are offered regarding the design of hydrogels that are required to function under physiological conditions over a long period of time.

2. Origin of Swelling and Strategy toward the Control of Swelling According to Flory et al., thermodynamics may describe the swelling or deswelling behavior observed after hydrogels are

Figure 1. Temporal change in the shape of hydrogels after the installation. A) Stage 1: hydrogels start to swell, and the swelling continues until the osmotic pressure becomes 0. B) Stage 2: the intentional or unintentional degradation of polymer chains induces further swelling. C) Stage 3: the complete disintegration of the polymer network takes place.

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Hiroyuki Kamata is a Research Fellow of the Japan Society for the Promotion of Science. He received his M.Eng. at The University of Tokyo in 2013, where he is currently pursuing his Ph.D. under the supervision of Prof. Ung-il Chung. His research interests include: the synthesis of functional polymers, the elucidation of the physical properties of stimuli-responsive hydrogels, and the exploration of their practical applications as structural biomaterials. Xiang Li is a Ph.D. student in the Tei/ Chung Laboratory at The University of Tokyo, where he is studying polymer physics and chemistry. He investigated the degradation behavior of polymer gels induced by the non-specific cleavage of network strands, and proposed a simple method to control the degradation behavior, without changing other physical properties. His research interests also include the dynamics of macromolecules in polymer gels. Takamasa Sakai received his Ph.D. in Engineering at the University of Tokyo. His doctoral research with Dr. Ryo Yoshida focused on analyzing the kinetics of the BelousovZhabotinsky reaction in polymer gels, and developing a novel selfoscillating gel system. He is currently an assistant professor at the University of Tokyo and investigating polymer gels with controlled network structure named Tetra-PEG gel. He focuses on the basics of polymer gels, including the mechanical properties and diffusion of substances in polymer gels. placed in an aqueous environment. Flory-Rehner theory is the most popular theory for predicting the equilibrium swelling state, where the total free energy is minimized with respect to the polymer volume fraction (φ).[19] For neutral hydrogels, the free energy is given by the sum of the mixing (Fmix) and




PROGRESS REPORT

One possible approach to suppress or control the swelling is to increase the elastic modulus (G); however, this approach itself is not sufficient to achieve the non-swelling property as is evident in Figure 2A. For instance, when we use the value of χ = 0.45, which is similar to that of poly(ethylene glycol) (PEG),[20] the Qe value surpasses 1 even in the case of a high elastic modulus (Qe ≈ 2 when G = 40 kPa). Note that G = 40 is higher than that of conventional synthetic polymer gels prepared as φ0 = 0.1.[21] Due to Figure 2. Theoretical prediction of the equilibrium swelling behavior of hydrogels. A) Q-depend- the above-mentioned problems, this approach ence of –∂Fmix/∂φ (dashed lines) and ∂Fel/∂φ (solid lines) with various χ (0.54, 0.52, 0.51, 0.50, is apparently not appropriate. To regulate the and 0.45) and G (1, 10, 25, and 40 kPa) values. Q = V/V0, where V is the volume in the equilibswelling behavior of hydrogels, therefore, rium swollen state, and V0 is the initial volume. Fmix and Fel are the mixing and elastic energies, respectively. φ is polymer volume fraction, χ is the Flory’s interaction parameter, and G is elastic the modification of χ parameters will also be required. Seeking polymer species with modulus. B) 3D plot of Q as functions of χ and G. χ ≈ 0.5 (θ condition),[19] although currently not commercially available, may result in the synthesis of hydrogels with a low degree of swelling, irreelastic (Fel) energies, and thus the equilibrium condition can be spective of G (Figure 2B). It is, however, not easy to synthesize written as: polymer gels with such χ values in an aqueous medium, as the ∂Fmix ∂Fel polymer components are more hydrophobic compared with (1) =0 + other water-soluble polymers (χ < 0.5); in that case, organic ∂φ ∂φ solvents may be required during the preparation steps at some point, which is not preferred for most of the potential biomedWhen a polymer gel contains elastically effective chains at ical applications, and undoubtedly precludes the applications of the concentration of ν (mol/m3), ∂Fmix/∂φ and ∂Fel/∂φ are given injectable hydrogels. as: With the current technologies available, the most feasible ∂Fmix approach toward the suppression of swelling is to use the phase 2 (2) = ( ln (1 − φ ) + φ + χφ ) transition behavior of thermoresponsive polymers. The dynamic ∂φ properties are often studied by employing such polymers as non-crosslinked polymers in aqueous solutions, rather than 1 ⎞ ⎛ cross-linked networks of hydrogels. Such polymers are water⎛ φ ⎞ ⎛ φ ⎞3 ∂Fel 1 (3) = νVs ⎜ − ⎜ ⎟ + ⎜ ⎟ ⎟ soluble at a low temperature, forming a transparent aqueous ∂φ ⎜ 2 ⎝ φ0 ⎠ ⎝ φ0 ⎠ ⎟ ⎠ ⎝ solution. Above a certain critical temperature (i.e., phase separation), on the one hand, the aqueous solution becomes turbid, and the solute starts to precipitate. The temperature, at which where φ0 is the polymer volume fraction in the preparation the phase separation occurs, is called lower critical solution temcondition, χ is the Flory’s interaction parameter, and Vs is the perature (LCST or Tc). Polymers with such properties include molar volume of solvent, respectively. Swelling ratio (Q), which is defined as the change in volume from the preparation condicopolymers of poly(ethylene glycol)/poly(propylene glycol), tion, is given by φ0/φ. ν is related with elastic modulus (G) as poly(glycidyl ethers), cellulose derivatives, poly(N-substitutedacrylamide).[22,23] Below LCST, polymer chains are hydrated due G = νRT (R: gas constant). Figure 2A shows Q-dependence of –∂Fmix/∂φ (dashed lines) to the strong interactions between water molecules and hydrophilic domains in the polymer chains. This hydration extends and ∂Fel/∂φ (solid lines) with various χ values calculated polymer chains, and let them take a random coil conformation. based on Equation (2) and (3), respectively. Here, φ0 was set Above LCST, on the other hand, dehydration takes place, and to be 0.1, which roughly corresponds to a hydrogel with the the polymer chains start to aggregate due to hydrophobic interwater content ≈90%. It was revealed from the figure that each actions. The phase transition behavior of aqueous solutions of combination of solid (G) and dotted (χ) lines has a unique poly(N-substituted-acrylamide) has been well disclosed among crossing point, at which the Q value corresponds to the equiothers, and the effects of substituent groups have been systemlibrium swelling ratio (Qe) (e.g., Qe ≈ 4 for a polymer gel with atically studied. The LCST of polymers can be adjusted via a χ = 0.45 and G = 10 kPa). Calculated crossing points were all copolymerization of monomers with different hydrophobicity; in the range Q > 1, except for the case with the highest χ value for example, copolymers synthesized from N-isopropylacryla(= 0.54), which suggests that swelling is a general tendency mide (NIPAAm), of which polymers (PNIPAAm) undergo of conventional gels. Figure 2B is a 3D plot that shows the phase transition at around 37 ºC, and relatively hydrophobic effects of χ and G on Q, with which again, we can safely say butyl methacrylate (BMA) exhibit LCSTs shifted to a lower that polymer gels with a low χ and G tend to swell, and vice temperature; conversely, copolymerizing NIPAAm and N,Nversa; in contrast, deswelling is observed in the condition with dimethylacrylamide (DMAA) is known to give those shifted to an extremely high χ or G.




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a higher temperature.[24] In a similar manner, a copolymerization of glycidyl ethers that bear different substituent groups with a proper copolymerization ratio produces polymers with a desired LCST.[22] As such, today we can design and synthesize a variety of thermoresponsive polymers. Hydrogels synthesized from LCST-type thermoresponsive polymers exhibit temperature-dependent volume changes (i.e., volume phase transition); at a low temperature, hydrogels swell since the component polymers exhibit a low χ parameter, while in contrast, at a high temperature, hydrogels deswell due to the dehydration of component polymer chains.[5] This dynamic phenomenon is typically used as an on-off switching function (i.e., transition between deswelling and swelling) in materials that require stimuli-responsiveness. It is also interesting to focus on the static “on” (deswollen) state, rather than the dynamic transition between the “on” and “off” states. Since previous studies showed that the equilibrium volume in the deswollen state depends on the composition of hydrogels,[25] one can say that hydrogels with properly combined hydrophilic and hydrophobic components result in the controlled degree of swelling. On top of that, exploiting this thermoresponsive mechanism allows for the design of injectable system; as LCST-type thermoresponsive polymers are water-soluble in the “off” state (i.e., hydrophilic at a low temperature), the aqueous polymer solution can be directly injected into the body. After being injected, the polymer network automatically undergoes the transition to the “on” state (i.e., hydrophobic at a physiological temperature), and continues to function as a switched-on device, demonstrating a controlled and constant degree of swelling.

3. Kinetics of Volume Phase Transition and Strategy for Accelerated Kinetics While the equilibrium “on” state of thermoresponsive hydrogels is of interest to control the swelling behavior, it is still critically important to focus on the dynamics of phase transition behavior. Tanaka et al. demonstrated that the cooperative diffusion of the polymer network governs the shrinking and swelling kinetics of polymer gels,[26] where the deformation process was well described in the linear response regime. Upon a sudden temperature jump across the critical temperature (Tc), however, the shrinking process of gels composed of thermoresponsive homopolymers is greatly decelerated, whereas the swelling process was well reproduced by the Tanaka-Fillmore theory.[27] This deceleration is explained today by the effect of phase separation that takes place not only on the surface (i.e., skin layer formation) but also in the inner core.[28] Conventional thermoresponsive gels, thus, have limited applicability in the actual situation. Recent experimental studies showed that the shrinking process can be accelerated by incorporating special mechanisms.[29,30] Among others, Hirotsu et al. succeeded in fabricating a new type of gels composed of both hydrophilic poly(acrylamide) (AAm) and thermoresponsive PNIPAAm to achieve improved shrinking kinetics.[30] This kind of network structure belongs to amphiphilic co-network (APCN),[31] which comprises two segmented domains with different hydrophilicity. The unique structure is known to contribute to the

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rapid shrinking kinetics during the volume phase transition of polymer gels; the continuous hydrophilic segments help water molecules pass through during the shrinking process. Most of the conventional APCN-type hydrogels, however, involve a free radical polymerization in the preparation steps, which inevitably introduces inhomogeneities to the resultant polymer network.[32] Not only do the inhomogeneities promote a heterogeneous phase separation retarding the shrinking kinetics, but also inhibit the correct understanding of the shrinking mechanism. Therefore, the development of polymer gels with a homogeneous network structure that do not involve heterogeneous phase separation has been of great interest both from the fundamental and technological points of view.

4. Design of Model APCN-Type Hydrogels To realize the precise control of swelling behavior of hydrogels, the polymer network should be simply modifiable, while excluding inhomogeneities. Historically, one of the most popular approaches to achieve controlled network structures is to use so-called “model network system,” where the polymer network is prepared from telechelic prepolymers that are able to connect to multi-functional crosslinkers.[33] However, inhomogeneities are observed even in such model network systems.[34] In 2008, we reported a new class of polymer gels, named tetra gels.[11] Tetra gels are prepared via an AB-type crosslinking of tetra-armed prepolymers, which are equipped with mutually reactive end-groups. We successfully demonstrated that tetra gels are free from inhomogeneities in the size range less than 200 nm, which suggests that this fabrication method has a potential to produce a near ideal polymer network.[35] Although the current generation of tetra gels primarily employs PEG as a backbone, tetra gels can be further functionalized by designing tetra-armed functional prepolymers.[36] The functionalization may be realized without losing the near ideal polymer network of conventional tetra gels. Similarly, we reported a model APCN-type hydrogel system, synthesized from tetra-armed hydrophilic PEG and thermoresponsive poly(ethyl glycidyl ether) (PEGE); the hydrogels comprise a simple APCN structure, where hydrophilic and thermoresponsive prepolymers are alternatingly aligned (TetraPEG-PEGE gels).[37] We employed the tetra gel backbone to fabricate the simple polymer network architecture. In this case, thermoresponsive segments incorporated in the gels shrink or swell in response to the change in temperature, inducing a macroscopic volume change (Figure 3A). Here, it is important to notice that hydrophilic domains always surround thermoresponsive segments. In APCN-type gels, hydrophilic polymer segments work as a pathway for water molecules during the shrinking process.[30] The similar effect is expected for the structure of Tetra-PEG-PEGE gels. The polymer network can be further modified through the proper selection of polymer unit ratio (rPEGE), without losing the perfection of the tetra gel backbone (Figure 3B).[25,37] Therefore, these rPEGE-tuned Tetra-PEGPEGE gels have a potential as ideal materials for quantitatively examining the true effect of the APCN structure on volume phase transition.




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increase in rPEGE (Figure 4B). This behavior stems from the phase transition nature of PEGE molecules.[22,38] Figure 4C shows the dependence of LCST on the PEGE concentration both for the aqueous solutions of PEGE and Tetra-PEG-PEGE gels. In aqueous PEGE solutions, Tc shifted from 14 °C to 6 °C with an increase in the PEGE concentration from 0.01% to 2%, which is the solubility limit of PEGE molecules in water. The PEGE concentration in Tetra-PEG-PEGE gels, on the other hand, was much higher than the solubility limit observed for the aqueous solution (≈2%), and the concentration was up to 5.4% in the case of rPEGE = 0.85. This increase in solubility suggests that hydrophilic segments exist near thermoresponsive segments, mutually affecting the hydration status.

4.2. Shrinking Kinetics Since Tetra-PEG-PEGE gels are categorized as an APCN gel, the improved shrinking Figure 3. Schematic diagrams of model APCN-type gels (Tetra-PEG-PEGE gels). A) Conceptual kinetics can be expected. The shrinking drawing of the volume phase transition mechanism in response to the change in temperature. Reproduced with permission.[37] Copyright 2012, RSC Publishing. B) Polymer network structure kinetics of Tetra-PEG-PEGE gels with difof Tetra-PEG-PEGE gels with various rPEGE values. Reproduced with permission.[25] Copyright ferent rPEGE values upon a sudden tempera2013, ACS Publications. ture jump was investigated. The change in temperature was set to be from 3 ºC to 40 ºC in order to cross Tc over all the samples. 4.1. Equilibrium Swelling Behavior Upon a temperature jump, Tetra-PEG-PEGE gels (rPEGE To characterize the unique network structure of Tetra-PEG≤ 0.75) shrunk homogeneously without inducing any surface PEGE gels, the equilibrium swelling behavior was well investiskin layer formation or apparent phase separation, and quickly gated. The swelling isotherms of rPEGE-tuned Tetra-PEG-PEGE reached the equilibrium state (Figure 5A). The photos during the shrinking process are displayed in Figure 5B. In the case gels were obtained in the temperature range from 3 °C to of rPEGE = 0.85, the gel shrunk forming bubbles on the surface; 40 °C (Figure 4A). At a low temperature (i.e., switched-off state), all the gels swelled, demonstrating the similar swelling this behavior is similar to the volume phase transition of conratio (i.e., Qe ≈ 3). With increasing temperature, the swelling ventional thermoresponsive polymer gels.[28] With regard to ratio dramatically changed at certain critical temperatures; a Tetra-PEG-PEGE gel with rPEGE = 1 (i.e., single-component this behavior corresponds to the phase transition of compoTetra-PEGE gel), the gel formed a dense skin layer on the surnent polymers. The rPEGE-dependent Tc was observed in the face immediately after the temperature jump, and the further shrinkage did not proceed over the scale of minutes. Tetra-PEG-PEGE gel system; Tc decreased linearly with an

Figure 4. Volume phase transition behavior of Tetra-PEG-PEGE gels. A) Equilibrium swelling ratio (V/V0) as a function of temperature, where V is the volume in the equilibrium swollen state at each temperature, and V0 is the volume of the as-prepared sample. The symbols represent the final molar ratio of PEGE (rPEGE); rPEGE = 0 (solid diamond), 0.15 (solid square), 0.25 (solid triangle), 0.35 (solid circle), 0.50 (open diamond), 0.65 (open square), 0.75 (open triangle), and 0.85 (open circle). B) Phase transition temperature (Tc) as a function of the final molar ratio of PEGE (rPEGE). C) Comparison of the phase transition temperature (Tc) of the aqueous solutions of Tetra-PEGE (solid circle), and Tetra-PEG-PEGE gels (open circle) as a function of the PEGE concentration. Reproduced with permission.[25] Copyright 2013, ACS Publications.




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Figure 5. Kinetic analysis of the shrinking behavior of Tetra-PEG-PEGE gels. A) Change in diameter upon a sudden temperature jump from 3 ºC to 40 ºC. The symbols represent the final molar ratio of PEGE (rPEGE); rPEGE = 0.15 (solid square), 0.25 (solid triangle), 0.35 (solid circle), 0.50 (open diamond), 0.65 (open square), 0.75 (open triangle), 0.85 (open circle), and 1.0 (solid diamond). B) Photos of the shrinking behavior of Tetra-PEGPEGE gels with different rPEGE values; rPEGE = 0.25, 0.50, 0.75, and 0.85. C) Variation of the relaxation of Tetra-PEG-PEGE gels (τsh) and D) variation of cooperative diffusion coefficient (Dsh) estimated from the relaxation time as a function of the final molar ratio of PEGE (rPEGE). Reproduced with permission.[25] Copyright 2013, ACS Publications.

The shrinking kinetics was further investigated quantitatively. Tanaka et al. demonstrated that the kinetics of polymer gels can be predicted from the diffusion process of a polymer network, which is characterized by cooperative diffusion coefficient (D). For spherical objects, the relationship can be written as follows: ∂u ∂ ⎧1 ⎡∂ ⎤⎫ = D ⎨ 2 ⎢ ( d 2u ) ⎥ ⎬ ∂t ∂d ⎩ d ⎣ ∂d ⎦⎭

(4)

where u is the normalized displacement, and d is the diameter of gels. Equation (4) theoretically holds regardless of the magnitude and the manner of deformation, and D is given as Mos/f, 6


where Mos and f are the osmotic modulus, and the friction coefficient of the polymer network, respectively. D is known to be identical to that obtained by dynamic light scattering (DLS). For homogeneous deformation, only the boundary condition on the surface is applied, and the following solution is obtained: dn =

d∞ − d ( t ) 6 t = 2 exp ⎛⎜ − ⎞⎟ ⎝ τ⎠ d ∞ − d0 π

( t >> τ )

(5)

where dn is the normalized size of gel, d(t) is the diameter of gels at a certain time (t), d0 is the diameter in the initial state, d∞ is the diameter in the equilibrium state, and τ is the characteristic time, respectively. Here, Equation (5) was applied to the




D=

3 d∞2 8 π 2τ

(6)

Using the relationship of Equation (6), D was plotted as a function of rPEGE (Figure 5D). The D values were almost constant against rPEGE (D ≈ 7 × 10−7 cm2 s−1), which well corresponded to that previously obtained for Tetra-PEG gels by DLS (i.e., rPEGE = 0, DDLS ≈ 6.5 × 10−7 cm2 s−1). This constant D irrespective of rPEGE, and the correspondence with DDLS strongly suggests that the cooperative diffusion coefficient in the equilibrium swollen state governs the shrinking kinetics of TetraPEG-PEGE gels, irrespective of the magnitude of the volume change. In stark contrast to the invariance of D in samples with rPEGE ≤ 0.75, the shrinking behavior of the gels with rPEGE ≥ 0.85 was completely different due to the skin layer formed and the heterogeneous deformation; therefore, the critical fraction of hydrophilic PEG segments (r*) that is required to form the path for leaking water molecules was estimated to be in the vicinity of 0.2. These results underpin the classical theory that hydrophilic domains help water permeate through toward the outer phase; in this case, hydrophilic PEG segments incorporated into the gels act as a pathway through which water can escape outside the gels during the shrinking process, and approximately one fourth may need to be hydrophilic segments to construct a percolated water path. These findings may help design future APCN-type hydrogels with improved volume phase transition behavior.

5. Design of Injectable Hydrogels with Controlled Swelling Behavior Demands for implantable hydrogels have been steadily increased in recent years. Sealants for surgical operations,[40] scaffolds for cells[41] are remarkable examples. To apply hydrogels to the body, one feasible way is to implant hydrogels that are already in the equilibrium swollen state, with which medical complications due to swelling may not be incurred. However, this complicates surgical procedures, and patients suffer from postoperative wound pain. For this reason, injectable system, which can minimize the burden on patients, has been demanded for minimally invasive surgery. Certain hydrogels can be treated as aqueous solutions during the preparation stage.[16] Although such hydrogels are found useful in some applications where mechanical robustness is not significantly


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Tetra-PEG-PEGE gel system because no heterogeneous deformation was observed in Tetra-PEG-PEGE gels (rPEGE ≤ 0.75). The characteristic time (τ) was estimated from the slope of the semi-log plots of dn against elapsed time. The values of τ are shown against rPEGE in Figure 5C. τ was a decreasing function of rPEGE; the larger deformation amount demonstrated the faster shrinking. The same analysis could not be conducted for samples with rPEGE ≥ 0.85, where the theoretical treatment is not appropriate due to the heterogeneous shrinking. Here, it should be noted that Equation (4) and (5) are formulated for spherical objects. For rod-shaped samples, τ is related to D as follows:[39]

important, the brittleness of conventional injectable gels still remains one of their critical problems. Unfortunately, it is not easy to incorporate special toughening mechanisms previously reported for so-called tough hydrogels, because they lack injectability or typically involve relatively complicated post-gelation preparation steps.[8–10] Messersmith et al. reported a novel toughening mechanism using a block copolymer composed of hydrophobic poly(propyrene oxide) (PPO) and hydrophilic PEG domains.[42] In this design, the shrinkage of core PPO domains balanced out the swelling of PEG domains, leading to the macroscopic “negative-swelling” of hydrogels at a high temperature. They also demonstrated that negatively swelling hydrogels exhibit improved mechanical strength even when the materials are in the equilibrium swollen state. However, their design may have limited applicability because the degree of swelling was uncontrollable due to the fixed ratio of PPO/PEG domains, and the in situ gelation property was not demonstrated. We designed and fabricated an injectable hydrogel system by combining tetra-armed hydrophilic PEG and thermoresponsive poly(ethyl glycidyl ether-co-methyl glycidyl ether) (Tetra-PEGP(EGE-co-MGE) hydrogels) (Figure 6).[43] Simple mixing of the aqueous solutions of component polymers gives hydrogels; this system does not involve any special techniques or knowledge about polymer chemistry, and more importantly users can freely design hydrogels with a desired degree of swelling for various purposes. The hydrogels with different thermoresponsive segment ratios (rth) can be easily prepared by the injection of the aqueous solutions of hydrophilic and thermoresponsive polymer units via a syringe; for example, when rth = 0, the hydrogel is composed only of hydrophilic segments. A completely alternating structure is formed when rth = 0.5. The gelation time after injection was controlled from seconds to hours by properly choosing the preparation condition. This strategy successfully led to the synthesis of hydrogels that can retain original mechanical properties even in an aqueous medium.

5.1. Swelling Properties The swelling behavior of Tetra-P(EGE-co-MGE) hydrogels is regulated by the thermoresponsive segment ratio (rth) (Figure 7A); all the hydrogels swelled in an aqueous environment at 10°C (Q ≈ 300%), while on the other hand, the hydrogels drastically changed their volume at around their Tc (≈25 °C, irrespective of rth). This constant Tc cannot be achieved in conventional APCN-type thermoresponsive hydrogel system, where gels are prepared via a random copolymerization of hydrophilic and thermoresponsive monomers; Tc inevitably increases with an increase in hydrophilicity of the polymer network, and eventually exceeds 37 °C,[44] which excludes the possibility of implantable applications. Since the hydrogel with rth = 0 is composed only of hydrophilic segments, the hydrogel behaves as a conventional hydrophilic gel. The swelling ratio of the hydrogel with rth = 0 was greater than 100% over the whole temperature range, which means that the hydrogel swells and alters its original shape. Therefore, it is not appropriate to apply such gels directly to an aqueous environment. To focus on the hydrogel with rth = 0.4, although the hydrogel swelled at a low



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Figure 6. Schematic of injectable hydrogels with controlled swelling behavior. Tetra-armed poly(ethylene glycol) with A) active ester end-groups and B) amino end-group. C) Tetra-armed poly(ethyl glycidyl ether-co-methyl glycidyl ether) with amino end-groups. D) Polymer network composed of hydrophilic (blue and green) and thermoresponsive (pink) polymer units, where rth represents the thermoresponsive segment ratio. E) Conceptual drawing of the control of swelling by introducing thermoresponsive segments that collapse at a physiological temperature. F) Gelation behavior after injection depending on the ionic concentration (I). Reproduced with permission.[43] Copyright 2014, American Association for the Advancement of Science.

temperature (e.g., 10 °C), the hydrogel recovered its original shape at 37 °C (Q ≈ 100%) (Figure 7B). Considering the fact that the average normal body temperature is around 37 °C, it is strongly suggested that Tetra-PEG-P(EGE-co-MGE) hydrogel with rth = 0.4 neither swells nor shrinks in the body. With the change in volume, the water content of Tetra-P(EGE-co-MGE) hydrogels also decreased; the degree of water release depended on rth (Figure 7C). It should be reminded here that the hydrogel with rth = 0.4 still retains a high amount of water at 37 °C (W0 ≈W ≈ 90%).

5.2. Mechanical Properties Theoretically, the mechanical properties of hydrogels are strongly affected by the degree of swelling.[19,45] To examine the physical properties of hydrogels with controlled swelling behavior, elongation tests on Tetra-P(EGE-co-MGE) hydrogels with different rth values were conducted. The representative stress–elongation curves showed that the maximum elongation ratio (λmax) decreased with a decrease in rth (Figure 8A). 8


This decrease in λmax can be explained by the following definition: λmax of polymer gels is defined as the ratio of the two lengths of network strands—the fully stretched and initial states.[45] Qualitatively, the pre-stretched network strands due to “swelling” lead to the decrease in λmax. The hydrogel with suppressed swelling (rth = 0.4) showed improved mechanical properties; the hydrogel did not fail even after being stretched more than sevenfold (Figure 8B), while a conventional hydrogel (rth = 0) in the equilibrium swollen state was easily torn off. Furthermore, the hydrogel with rth = 0.4 showed practically no hysteresis, at least when stretched less than fourfold, which indicates that the stretching did not break the covalent bonds of the polymer network (Figure 8C). This reversible feature is vital for applications where the materials are required to tolerate a continual mechanical load. Notably, the hydrogel with rth = 0.4 endured a compressive stress of up to 60 MPa even though the hydrogel was in its equilibrium swollen state (Figure 8D), which is comparable to so-called tough hydrogels previously reported.[9–11,45] In stark contrast, the swollen hydrogel (rth = 0) fractured at a strain of ≈80%, showing a maximum stress of ca. 0.4 MPa.




PROGRESS REPORT Figure 7. Swelling behavior of Tetra-P(EGE-co-MGE) hydrogels. A) Swelling ratio (Q) as a function of temperature. The symbols represent rth; rth = 0 (open square), 0.1 (open triangle), 0.2 (open circle), 0.3 (asterisk), 0.4 (solid circle), and 0.5 (solid triangle). Q = V/V0 × 100, where V is the volume of the samples in the equilibrium-swollen state at each temperature and V0 is the initial volume of the samples. The dotted line is a guide to the eye for 100%. B) Photos of Tetra-P(EGE-co-MGE) hydrogels that exhibit different swelling degrees depending on rth; as-prepared samples (top) and samples equilibrated in D-PBS at 37 °C (bottom). The transparent hydrogels were colored only for visibility. C) The initial water content (W0) (open inverted triangles) and the equilibrium water content (W) of Tetra-P(EGE-co-MGE) hydrogels at 10 ºC (solid squares) and 40 ºC (solid triangles). Reproduced with permission.[43] Copyright 2014, American Association for the Advancement of Science.

Figure 8. Results of mechanical tests on hydrogels with controlled swelling properties. A) Representative stress-elongation curves of the hydrogels equilibrated in D-PBS at 37 ºC. Only for rth = 0 (as-prepared), the sample was measured immediately after being taken from the mold. The color represents rth. B) Photos of the hydrogels during the elongation tests. C) Results of hysteresis test on Tetra-P(EGE-co-MGE) hydrogels. The hydrogel with rth = 0.4 was tested after being equilibrated in D-PBS at 37 °C. The color represents the order of elongation test. D) Result of compression test on Tetra-P(EGE-co-MGE) hydrogels. The hydrogel with rth = 0.4 was tested after being equilibrated in D-PBS at 37 °C. σ and ε represent stress and strain, respectively. Reproduced with permission.[43] Copyright 2014, American Association for the Advancement of Science.




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Figure 9. Schematic illustration of hydrogels with degradable sites. A) three different prepolymers: tetra-armed poly(ethylene glycol) with amino endgroup (TAPEG), active ester end-groups (TCPEG), and active ester end-groups that bear ester bonds (TGPEG).The chemical structure of the end groups of B) TCPEG and C) TGPEG. D) Time course of swelling ratio (Q) of Tetra-PEG gel with degradable units. rdeg = 0 (solid circle), 0.63 (solid square), 0.69 (solid triangle), 0.81 (open circle), 0.88 (open square), 1.0 (open triangle). E) Time course of cycle rank density (ξ) of gel samples with different degradable units. rdeg = 0 (solid circle), 0.63 (solid square), 0.69 (solid triangle), 0.81 (open circle), 0.88 (open square), 1.0 (open triangle). The solid curves show the fitting curves of our model. F) Disintegration time (tdeg) as a function of rdeg. Reproduced with permission.[36] Copyright 2013, ACS Publications.

6. Precise Control of Degradation Behavior of Hydrogels The degradation of hydrogels has been actively utilized for specific applications.[46] In drug delivery systems, for example, the drug release rates may be adjusted by the controlled degradation. For scaffolding materials used in the regenerative medicine, hydrogels need to provide the space to regenerating tissues, where the synchronized degradation must be accompanied. As such, the precise control of the degradation behavior of hydrogels is of great importance in the biomaterials field. Numerous studies have been conducted on hydrogels that are equipped with controlled degradation behavior.[47] One of the most popular methods to control the degradation behavior is to tune the polymer volume fraction and/or the crosslinking density of the polymer network. This kind of methods, however, inevitably alter other important physical parameters, such as water content, elastic modulus, deformability, and so on.

6.1. Design of Hydrogels with Controlled Degradation Behavior In 2011, we proposed a unique molecular design to independently control the degradation behavior without affecting other

10


physical parameters.[36] Our strategy involves the quantitative control of cleavable sites while maintaining the polymer volume fraction and crosslinks density. We employed the tetra gel backbone, where all the cross-links are formed as amide bonds. To introduce cleavable bonds in a controlled fashion, we introduced a new tetra-armed prepolymer that bears ester bonds, which are subject to hydrolysis in an aqueous environment, in their terminal groups (TGPEG, Figure 9A). The only difference from the first generation of tetra gels is, therefore, that the resultant polymer network comes with or without cleavable bonds located near the amide junction points (Figure 9B). In this study, the ratio of cleavable unit is defined as rdeg; for example, gels with rdeg = 1 contain ester bonds at all the junction points, while on the other hand, rdeg = 0 indicates none of polymer chains contains cleavable sites.

6.2. Modeling and Analysis We introduced a new theoretical model for the degradation behavior of hydrogels,[36] in which the polymer network structure was treated using the tree-like theory,[48] a theoretical model that relates the cycle rank density (ξ) of the network with the connectivity (p) of the network strands. For the polymer




1 1 ⎞ ⎞⎛ ⎛ ⎛ 1 3⎞ 2 3 ⎛ 1 3⎞ 2 1 ξ =U ⎜ + ⎜ − ⎟ ⎟ ⎜ − ⎜ − ⎟ ⎟ ⎜ 2 ⎝ p 4⎠ ⎟ ⎜ 2 ⎝ p 4⎠ ⎟ ⎠ ⎠⎝ ⎝

3

(7)

where U is the molar concentration of the prepolymers in the as-prepared state. Assuming that the hydrolysis of ester groups (i.e., degradation sites) follow pseudo-first order kinetics, probabilities that ester groups exist after a certain period of time (pester) are given as: pester = exp ( −kdeg t )

(8)

where kdeg is the degradation rate constant of ester groups. For gels with rdeg, the probability or connectivity of the network strands that survive after a certain period of time is given as:

(

p = p0 (1 − rdeg ) + rdeg exp ( −kdeg t )

)

(9)

where p0 is the initial connectivity of the network strands in gels. By combining Equation (7) and (9), ξ can be related with rdeg and t as: 1 ⎡ ⎤ 2 ⎢1 ⎛ 1 3⎞ ⎥ ξ (t ) = U ⎢ + ⎜ − ⎟ ⎥ ⎢ 2 ⎜⎝ p0 (1 − rdeg ) + rdeg exp ( −kdeg t ) 4 ⎟⎠ ⎥ ⎣ ⎦ 1 3 ⎡ ⎤ 2 ⎢3 ⎛ 1 3⎞ ⎥ − − ⎜ ⎟ ⎢ ⎥ ⎢ 2 ⎜⎝ p0 (1 − rdeg ) + rdeg exp ( −kdeg t ) 4 ⎟⎠ ⎥ ⎣ ⎦

(

(

)

)

(10)

To assume that hydrogels disintegrate at a critical connectivity (pc), the time of disintegration (tdeg) is given as following according to Equation (9): tdeg =

1 kdeg

ln

rdeg pc − (1 − rdeg ) p0

(11)

To validate the theory, the influence of rdeg on elastic modulus (G) was evaluated. For the quantitative evaluation, the cycle rank density (ξ0) in the as-prepared state was calculated from G, based on the phantom network model with the treelike approximation (Equation (7)). ξ0 was almost constant, regardless of rdeg, which suggests that the replacement of the tetra gel backbone with degradable TGPEG units did not affect the architecture of tetra gels. The values of ξ0 corresponded to those at a reaction conversion (p0) of 0.9. Degradation tests for the gels were performed in phosphate buffer (pH7.4) at 37 °C. Gel specimens swelled over time, resulting in the bulk degradation (Figure 9D), and the degree of swelling was pronounced when rdeg was increased. No swelling was observed in the case of rdeg = 0, which indicates that hydrogels without cleavable linkers do not degrade in this experimental condition. The time evolution of ξ was then estimated based on the Flory-Rehner theory with the phantom network model (Equation (18)) (Figure 9E):[19]


PROGRESS REPORT

network prepared by tetra-armed prepolymers, the relationship between ξ and p can be written as:

1

⎛φ ⎞3 −V ⎜ e ⎟ ξ = ln (1 − φe ) + φe + χφe2 ⎝ φ0 ⎠

(12)

where V is the molar volume of solvent, χ is the Flory interaction parameter, φ0 and φe are the polymer volume fraction in the as-prepared state and in the equilibrium swollen state, respectively (Qe = φ0/φe). The values of ξ of gels with rdeg = 0 were almost constant in the tested time range, while, in contrast, ξ of the gels with rdeg > 0 decreased over time. The decreasing rate was accelerated with an increase in rdeg. A fitting analysis of ξ with Equation (10) resulted in a good agreement. Here, it should be noted that only kdeg was used as a fitting parameter, while p0 and rdeg are fixed. kdeg estimated was constant against rdeg, which underpins the validity of the proposed model. The proposed model was further extended to develop the expression for the prediction of the relationship between the time of disintegration (tdeg) and rdeg. Based on tdeg in Figure 9D as well as the time evolution of ξ in Figure 9E, we estimated the critical connectivity (pc) at the disintegration point with Equation (18). pc was constant (=0.46), regardless of rdeg, which suggests that all the gel specimens disintegrated at a universal value of pc, regardless of rdeg. This value of pc is similar to that predicted for a 3D diamond lattice under the percolation model (pc = 0.43),[17] which is another factor that can support the validity of the proposed model and the existence of the homogeneous network. We finally confirmed the consistency of the prediction of our expression (Equation (18)) (Figure 9F).

7. Degradation of Hydrogels by the Non-Specific Cleavage of Network Strands Hydrogels are considered as candidate materials for the design of artificial locomotor systems. Under actual physiological conditions, hydrogels inevitably degrade over time by various kinds of stress, which includes oxidative and mechanical stress. This behavior is distinct from that of natural locomotor organs, of which the structure is dynamically maintained by the continuous breakage and reconstruction. To predict the stability of hydrogels on a long-term basis, we need to consider the situation that each molecular segment in the polymer network is cleaved in a non-specific manner. The correct understanding of this non-specific degradation is of particular importance in the design of hydrogels, because hydrogels are the open system; the openness may cause the bulk degradation of hydrogels, which is distinct from the surface degradation observed in other closed materials, such as metals and ceramics. In the bulk degradation process, the degradation does not occur only on the surface, but occurs in the entire body, resulting in the drastic deterioration in the physical properties. Despite the importance, there are few studies regarding the non-specific degradation behavior of hydrogels.

7.1. Modeling of Degradation Behavior We conducted the first systematic analysis of the non-specific degradation of hydrogels,[17] where we proposed a theoretical



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Figure 10. Schematic illustration of the degradation tests on hydrogels with the tetra gel backbone. Reproduced with permission.[17] Copyright 2014, ACS Publications.

model that can describe the non-specific degradation behavior of hydrogels, and experimentally verified the model using hydrogels with the tetra gel backbone as a model system. The network strands of the hydrogels are formed with three different chemical bonds (Figure 10): carbon-carbon (R1CCR2), ether (R1-O-R2), and amide bonds (R1-CONH-R2). In an aqueous hydrogen peroxide (H2O2) solution, ether and amide bonds may be preferentially cleaved by oxidation and/or hydrolysis, when compared with carbon-carbon bonds. Our analysis assumed that the cleavage reaction of ether and amide bonds follows pseudo-first order kinetics. We defined the possibilities that an ether bond and an amide bond survive after a period of time (pether and pamide, respectively) as: pether = exp ( −kether t )

(13)

pamide = exp ( −kamide t )

(14)

where kether and kamide are degradation rate constants of ether and amide bonds, respectively. To consider the chemical structure of the hydrogels, there are N ether bonds and one amide bond in each network strand. Thus, the probability that a network strand survives after a certain period of time (p) can be expressed as:

(

)

p = p0 ( pether ) pamide = p0 exp [ − (kamide + Nkether ) t ] N

(15)

where p0 is the connectivity of prepolymers in the as-prepared state. We also assumed that the polymer network of the hydrogels can be described by the tree-like theory, which is a model that relates ξ and p.[49] By substituting p (Equation (15)) into the treelike theory (Equation (7)), ξ may be expressed as a function of t as: 1 ⎡ ⎤ 1 ⎛ 1 3⎞ 2 ⎥ ⎢ ξ (t ) = U ⎢ + ⎜ − ⎟ ⎥ 2 ⎝ p0 exp ( −kapp t ) 4 ⎠ ⎢⎣ ⎥⎦ 1 3 ⎡ ⎤ 1 3⎞ 2 ⎥ ⎢3 − ⎛ − ⎢ 2 ⎜⎝ p0 exp ( −kapp t ) 4 ⎟⎠ ⎥ ⎣⎢ ⎦⎥

12


(16)

kapp = kamide + Nkether

(17)

where U is the molar concentration of the prepolymers in the as-prepared state, and kapp is the apparent degradation rate constant of a network strand. U can be calculated form the prepolymer concentration in the as-prepared state. p0 and kapp are free fitting parameters.

7.2. Accelerated Degradation Testing Based on the prediction from Equation (18), the apparent degradation may be pronounced with an increase in N. To validate the prediction, we prepared hydrogels with a variety of polymerization degrees of network strands (N = 114, 228, and 456), and performed the accelerated degradation testing in an aqueous H2O2 solution (30 wt%) at certain temperatures (60, 64, 70 and 80 °C). Just after the immersion to the aqueous H2O2 solution, gel specimens started to swell, and reached the equilibrium in 30 min. After experiencing a several minutes of plateau region, gel specimens continued to swell, showing the bulk degradation behavior by the non-specific cleavage. The Flory interaction parameters (χ) were estimated from the equilibrium swelling ratio (Qe) obtained during the first 30 mins using the Flory-Rehner equation for the phantom network model (Equation (12)). Figure 11A shows χ of hydrogels with different N as a function of the absolute temperature (T). χ increased with an increase in T, and χ decreased with an increase N. Both tendencies are well known characteristic of PEG.[20,49] The obtained values of χ were used as universal parameters for each experimental condition. According to Equation (12), the values of Qe were converted to the cycle rank density (ξ). ξ decreased over time (t), reflecting the swelling. The time course of ξ was analyzed using Equation (16) to estimate the apparent degradation rate (kapp). The fitting analysis worked well in a wide range of time for all the experimental conditions (Figure 11B). The p0 obtained from the fitting well corresponded to the p0 measured directly by infrared spectroscopy with the error of ± 2%. Figure 11C shows the estimated kapp as a function of N. kapp increased linearly with an increase in N. The linear relationship




PROGRESS REPORT Figure 11. A) The Flory interaction parameter (χ) of the different gel samples as a function of T. B) Time course of cycle rank density of hydrogels with the tetra gel backbone during the degradation in aqueous H2O2 solution (30 wt%) at 60 °C. The symbols represent the molecular weight of prepolymers: (open circle) 10k, (solid square) 20k, and (solid triangle) 40k. The gels made of prepolymers with molecular weights of 10k, 20k, and 40k have the polymerization degree (N) of network strands as 113, 228, and 456, respectively. The solid curves represent the fitting curves with Equation (16). C) kapp as a function of N at different T. The dotted lines illustrate the fitting curves with Equation (17). D) Semi-logarithmic plot of kether/T and kamide/T as a function of 1/T. The dotted lines show the fitting curves of k/T = A exp(–ΔH/RT). Reproduced with permission.[17] Copyright 2014, ACS Publications.

between kapp and N well agreed with the prediction (Equation (17)). Based on the prediction, we estimated kamide and kether from the slope of the extrapolated lines. Theoretically, the temperature dependence of kamide and kether should obey the transition state theory: ⎛ ΔS * ⎞ ⎛ ΔH * ⎞ k ≈ T exp ⎜ exp ⎜ − ⎟ ⎝ R ⎠ ⎝ RT ⎟⎠

(18)

where k is the reaction rate constant, ΔS* is the activation entropy, ΔH* is the activation enthalpy and R is the gas constant (8.314 J mol−1 K−1).[50] Figure 11D shows the relationship between ln(kamide/T) as a function of 1/T. Linear relationships were observed both between log(kamide/T) and 1/T, and between logkether and 1/T, which indicates that both reaction rate constants follow the transition state theory. The estimated values of ΔH* of amide and ether bonds were comparable with each other, and to those estimated in the previous studies on similar experimental conditions.[51] The values of A (∼ exp(ΔS*/R)) of amide bonds were approximately 2 orders of magnitude larger than those of ether bonds. We provide the reason for this large discrepancy as the difference in the number of cleavage pathways of ether and amide bonds; ether bonds only degrade by the oxidation by H2O2, whereas amide bonds suffer from both the oxidation by H2O2 and hydrolysis from H2O. Based on our experimental results on ΔH* and ΔS*, one can estimate kapp of hydrogels with the tetra


gel backbone at 37 °C, which is roughly one order of magnitude smaller than that at 60 °C. This suggests that the disintegration time is around 1–3 weeks at 37 °C in a 30 wt% H2O2 aqueous solution. Although the radical species generated in body is expected to be much smaller than that used in this study, we cannot ignore the effect of degradation in materials used in a time range of a year.

8. Current Challenges and Future Research Directions Indeed, it has been demonstrated that the swelling behavior of hydrogels can be controlled by the incorporation of thermoresponsive segments, where the swelling at an early stage is well suppressed (Figure 1A). This allows for the use of hydrogels in an aqueous medium, including physiological environments, for a certain period. Yet, a next challenge will be to establish countermeasures for degradation-induced swelling (Figure 1B). In some biomedical applications, hydrogels are required to degrade with a programed mechanism. Otherwise, the polymer network is exposed to oxidative or mechanical stress, resulting in non-specific degradation. Recent experimental studies showed that hydrogels degrade faster than predicted for the simple chemical structure of the backbone;[17] the apparent degradation rate constant (kapp) is the sum of the degradation rate constants of all the chemical bonds that consist of



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the entire network strand (i.e., kapp = Nkm, where km is written as the sum of the degradation rate constants of all the bonds in monomeric units, and N is the degree of polymerization of a network strand). In typical hydrogels, N can be up to 100, which makes it difficult to maintain the mechanical properties of hydrogels on a long-term basis. As such, polymeric materials are, in essential, susceptible to non-specific degradation, which steadily leads to degradation-induced swelling. Thus far, this aspect, although critically important, has rarely been discussed in the biomaterials field. For instance, hydrogels that comprise polyethers, such as PEG, which is a well-known biocompatible polymer, are inherently chemically unstable when exposed to reactive oxygen species;[17] thus, it is not appropriate to expect a long lifetime from polyethers under physiological conditions. A future approach to overcome degradation-induced swelling will be to seek or synthesize polymers with an optimal χ value by modifying the primary chemical structure of polymer chains. For example, non-swelling polymers will produce gels with more or less suppressed swelling behavior (Figure 2). Further, it could be speculated that degradation-induced swelling may also be suppressed or even negligible, as swelling is suppressed at a molecular level. This also allows for the control of swelling without being affected by elastic properties, which should be tailored for each purpose. While it is not directly mentioned in this report, initial concentration (φ0) is also a key parameter to determine the degree of swelling. Osmotic pressure at an early stage becomes low when hydrogels come with a low φ0; therefore the swelling ratio will be minimized. Although it seems effective, one crucial disadvantage of this strategy is that hydrogels with a high elastic modulus cannot be obtained since, in general, low concentrations contradict high elastic moduli. Still, it may find applications for soft tissues with an elastic modulus of less than ten kilopascals, such as eye balls[52] and vocal folds,[53] and so forth. Nonetheless, it is still challenging to synthesize gels at a low initial concentration, where cross-linking does not effectively work. To achieve such gels, new synthetic strategies have to be explored; the cross-linking of dendritic prepolymers or the use of pre-grown clusters may be useful.

The importance of the control of swelling has been gradually recognized, and some practical methods have been already introduced in the polymeric materials field. Swelling at an early stage after installation is well controlled with the current technology. The long-term stability of hydrogels in the body has yet to be further investigated. Degradation-induced swelling is particularly concerned when used under mechanical stress on a long-term basis. The synthesis of polymer chains with a controlled χ parameter will be a next challenge, with which we can synthesize hydrogels that stably function under physiological conditions on a mid- and long-term basis. This technology can open up significantly more opportunities in the development of biomaterials used in a wet environment, such as wound dressings, scaffolding materials, and drug reservoirs.


This work was supported by the Japan Society for the Promotion of Science (JSPS) through the Grants-in-Aid for Scientific Research, the Center for Medical System Innovation (CMSI), the Graduate Program for Leaders in Life Innovation (GPLLI), the International Core Research Center for Nanobio, and the Funding Program for World-Leading Innovative R&D on Science and Technology (FIRST program); the Ministry of Education, Culture, Sports, Science, and Technology in Japan (MEXT) through the Center for NanoBio Integration (CNBI); the Japan Science and Technology Agency (JST) through the S-innovation program; and Grant-in-Aids for Scientific Research from the Ministry of Education, Culture, Sports, Science and Technology (no. 23700555 to TS and no. 24240069 to UC). Received: January 30, 2015 Revised: March 20, 2015 Published online:

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