From Biomaterial, Biomimetic, and Polymer to Biodegradable and

Chapter 1

Downloaded by 80.82.77.83 on November 1, 2017 | http://pubs.acs.org Publication Date (Web): October 25, 2017 | doi: 10.1021/bk-2017-1253.ch001

From Biomaterial, Biomimetic, and Polymer to Biodegradable and Biocompatible Liquid Crystal Elastomer Cell Scaffolds M. Prévôt1 and E. Hegmann1,2,3,* 1Liquid

Crystal Institute, Kent State University, Kent, Ohio 44242-0001, United States 2Chemical Physics Interdisciplinary Program, Kent State University, Kent, Ohio 44242-0001, United States 3Department of Biological Sciences, Kent State University, Kent, Ohio 44242-0001, United States *E-mail: [email protected]

3D scaffolds are no longer simply physical templates for cell growth and tissue formation; they also have to provide chemical, biomolecular, mechanical and geometrical signals to cells. Liquid crystals (LCs) can be of significant importance in tissue engineering because they are able to report anisotropic growth of expanding cell lines back to the observer with an easily discernable optical response such as a change in birefringence or via the alignment of the LC molecules. The real challenge is to design materials that can direct or guide the behavior of biological materials. After a brief survey of several materials classes intensively investigated for the use as 3D cell scaffolds, we summarize recent research on a particular class of LC materials, liquid crystal elastomers (LCEs), that can serve as unique, longitudinal and multi-responsive cell scaffolds suitable for cell attachment, cell proliferation and cell alignment. Several types of biocompatible, biodegradable LCE scaffolds with specially engineered porous architectures will be introduced and their advantages discussed.

© 2017 American Chemical Society Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


Introduction Tissue, in a multicellular organism, is defined as an aggregate of similar cells forming a definite kind of structural material with a specific function, whereas an organ is the product of several similar tissues grouping together. There are in principle four types of tissue: connective, muscle, nervous, and epithelial tissue. Connective tissue binds tissues and organs in the body, providing cohesion and internal support. Connective tissue is mainly composed of fibrous tissues formed from non-living material secreted by living cells called the extracellular matrix (1) (ECM). The ECM is composed of two classes of macromolecules: fibrillar proteins (e.g., collagen, elastin) (2), bestowing mechanical and bioactive properties to the matrix (3), and glycosaminoglycans, that correspond to long unbranched polysaccharides (e.g., hyaluronic acid) (4), as shown schematically in Figure 1.

Figure 1. Description of extracellular matrix. Reproduced from reference (1). Copyright 2016, ACS Biomater. Sci. & Eng. Lett. Tissue engineering consists in the formation of functional substitutes for the therapeutic reconstruction of damaged or diseased tissues by the careful and controlled stimulation of selected target cells through a systematic combination of molecular and mechanical signals (5). Cells express functional responses by receptors on their surfaces that are mediated by the ECM. These responses convey developmental decisions, cell migration, cell maturation as well as differentiation, cell survival, tissue homeostasis and, in some cases, tumor cell invasion (6). Tissues are constantly in stress and must be resilient enough to deform reversibly without damage and still maintain their functions. Tissues must acquire specific mechanical properties to create an appropriate cell environment for the development of the ECM. The process of engineering tissue involves the design and use of a support that maintains tissue contour, particularly in the form of a 3D scaffold, implanted at the defective site. An ideal scaffold should then satisfy certain specific requirements (7, 8); it should: i.

be biocompatible (non-toxic) and should provide a framework for cells to attach and proliferate. Biocompatibility refers to the ability to co4 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


exist with human tissues without causing an inflammatory or a rejection response (9); ii. exhibit, during and after in vitro tissue culture tests, mechanical properties that can mimic native environments especially elasticity found in dynamic and structural tissues; iii. have a porous 3D architecture to allow cell growth, vascularization and transport of nutrients between the cells seeded within the matrix and the surroundings (10). As the properties of the ECM are characteristic of each tissue, so then engineering scaffolds must be designed to emulate the molecular and structural properties of the native ECM (11–13). Moreover, a well interconnected, highly porous structure is necessary to allow high cell seeding density and unimpeded tissue growth (14); and finally iv. degrade at a rate that matches the regeneration of new tissue and into non-toxic products that can easily be resorbed or excreted first by the cells, and finally by the body. Meaning that cells inside the scaffold can merge with the tissue at the implantation site and induce new tissue to infiltrate, while the scaffold material itself is gradually degraded in vivo (15). Biocompatibility Biocompatibility depends of the capability of a particular material to co-exist within body tissues without causing any considerable harm. Williams et al. defined biocompatibility as the “ability of a material to perform with an appropriate host response in a specific situation” (9). The biocompatibility of biomaterials is classified according to their ability to induce damage as described in Table 1. Thus, materials are expected to be non-toxic, non-immunogenic, non-thrombogenic, non-carcinogenic, and a non-irritant.

Table 1. Different types of biocompatibility problems (16). Name

Cause induced

Cytotoxicity

Cell or tissue death

Carcinogenicity

Cancer formation

Mutagenicity

Damage genes

Pyrogenicity, allergenicity

Immune response

Thrombogenicity

Blood clotting

According to Williams et al., biodegradability is not limited to immunological, toxic, or foreign body responses. Materials should be expected to passively allow or actively produce palpable beneficial effects in that particular host (9). Tissue constructs continuously interact with the body through the healing and 5 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


cellular regeneration process as well as during scaffold degradation. Because cell–matrix interactions influence a variety of important processes, it is essential that biomaterial scaffolds facilitate cell adhesion to the biomaterial surfaces to prevent apoptosis in cells. Cell-matrix anchorage is mediated through membrane proteins. The tensile strength of the ECM is a result of collagen; and fibronectin is responsible for adhesion of cells to the matrix (16). However, including both chemical and topographical characteristics, surface properties also control and affect cellular adhesion (17) as for example, hydrophilicity, geometrical features of the porous structure (pore size, pore size distribution), micro-architecture, heterogeneity, (an)isotropy, and interconnectivity of the porosity within the scaffold (9). Performance of biomaterials can be enhanced by surface modification such as, inclusion of short peptide motifs like the well-known Arg-Gly-Asp (RDG, tripeptide composed of L-arginine, glycine, and L-aspartic acid, a recognition system for cell adhesion) (18).

Mechanical Properties Before analyzing some functional mechanical attributes of materials, we need to clarify the definitions of stiffness and strength. Stiffness measures the ability of an elastic solid material to resist deformation to an applied stress. It is represented by the elastic modulus, also named Young’s modulus (YM). Strength corresponds to the maximal amount of tensile stress before failure. Among mechanical properties used to fully characterize tissue, toughness, extensibility, spring efficiency, durability, and spring capacity should all be considered (19). Table 2 summarizes functional attributes of materials and give the corresponding property and its unit. It should also be considered that growing cells cannot be exposed to more than 30% tissue extensibility without risk of damaging the cellular membrane. Stiffness and strength become then the main parameters to study scaffold materials. The mechanical properties of the ECM vary from tissue to tissue. Table 3 gives the YM of common soft body tissues. McKee et al. (20) have shown that the YM values depend on the method used to measure them. On the one hand, the YM can be obtained by applying a force to a section of tissue and measuring the change of length or strain (Figure 2 right). An alternative method involves controlled poking by indenters, including atomic force microscopy tips (Figure 2 left). The YM obtained by indentation measurements are usually quite different from those acquired by tensile measurements, (i.e., in the bulk of material). For example, liver and kidney tissues possess a much lower tensile YM compared to muscle tissues (10 MPa vs. 480 MPa) but show higher indentation YM (190 kPa vs. 7kPa) (20). Thus, to properly interpret cellular responses to biophysical stiffness there is a need to characterize both the indentation and tensile values of the YM, especially for non-homogeneous tissues (20). Table 3 summarizes indentation and tensile measurements of the moduli of common soft body tissues. 6 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


Table 2. Some functional attributes of materials and the material properties and units used to quantify these attributes. (taken from Gosline et al. (19)) Functional attribute

Material property

Units

Stiffness

Modulus of elasticity, Eint

Nm−2 (Pa)

Strength

Stress at fracture, σmax

Nm−2 (Pa)

Toughness

Energy to break work of fracture

Jm−3, Jm−2

Extensibility

Strain at failure, εmax

No units

Spring efficiency

Resilience

%

Durability

Fatigue lifetime

s to failure or cycles to failure

Spring capacity

Energy storage capacity, Wout

Jkg-1

Table 3. Indentation vs. tensile measurements of Young’s moduli. Data taken from McKee et al. (20) Tissue

Indentation (kPa)

Tensile (MPa)

Skin

~85

~30

Liver & kidney

~190

~10

Spinal cord & gray matter

~3

~2

Muscle

~7

~480

Tendon

No values

~560

Breast tissue

~8

No values

Artery & vein

~125

~2

Sclera

No values

~2.7

Cornea

~29

~3.0

Figure 2. Illustration of the two methods used to measure the anisotropy of the YM: indentation (left) and tensile (right).

Three-Dimensional (3D) Scaffolds Scaffolds play a pivotal role in tissue engineering and a porous architecture within these scaffold must allow for sufficient transport of oxygen, nutrients, metabolites, and cellular signals (among other regulatory factors) (21). Different 7 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


types of scaffolds have been studied by many groups including hydrogel-based scaffolds, microsphere-based scaffolds, fibrous scaffolds, and polymeric porous scaffolds. For these main scaffolds, fabrication techniques include a variety of methods such as salt leaching, gas foaming, phase separation, electrospinning, using of free-form, and lithography approach. A comparison of these fabrication techniques has been presented in a review-article by Seunarine et al. (22). Hydrogels are crosslinked hydrophilic polymers capable of swelling and retaining a large amount of water within their 3D network without dissolution. They possess a degree of flexibility very similar to natural tissue due to their water content, allowing minimally invasive procedures, and exhibit excellent biocompatibility (23). Synthetic polymers used in tissue engineering include especially poly(ethylene glycol) (PEG) (24–27), poly(vinyl alcohol) (PVA), and polyacrylates. Hydrogels acts as bulk 3D gel framework to which cells are adhered or suspended (28). Hydrogels can also be designed to incorporate bioactive agents (growth factors, proteins), however, they commonly do not guide cells to alignment and differentiation by applied external stimulus (29, 30). Microspheres can be fabricated using a variety of different biodegradable polymers and with surface porosity to improve diffusion of nutrients and oxygen (31). Microspheres as scaffold building blocks offer several benefits: (i) their fabrication, morphology, and physicochemical characteristics are simple to control and (ii) they allow simple control of the release kinetics of encapsulated factors (32). A 3D nano-fibrous scaffold can be generated by electrospinning, molecular self-assembly, or thermally induced phase separation (TIPS) of a polymer mixture. Phase separation has shown high potential to meet the needs of 3D tissue regeneration due to its ability to incorporate any pore shape and size or any overall 3D geometry (33). The high surface-area-to-volume ratio of the nanofibers combined with their microporous structure favors cell adhesion, proliferation, migration, and differentiation. Nano-fiber technology, however, must place the cells within the nano-fibrillar structure and form the porous network in situ without cellular damage, resulting in difficult processability (34). 3D polymeric porous scaffolds have been widely used for tissue engineering. The typical scaffold design includes sponge or foam morphologies. These designs promote a homogeneous, interconnected pore networks. These porous scaffolds can be manufactured with specific pore sizes, porosity, and surface-area-to-volume ratios (17). Moreover, they offer versatility of the employed chemistry. Later in this chapter, we will focus only on porous scaffold. Scaffolds, in general, should be morphological and architecturally similar to the host environment. They should also promote vascularization; permitting new blood vessels to grow implies appropriate porosity and pore size. Additionally, a high surface-area-to-volume ratio in interconnected porous scaffolds favors cell attachment and proliferation. Thus, scaffolds must possess a highly porous structure with an open and fully interconnected geometry (17) with high structure porosity (28). If the pores are too small, pore occlusion by the cells can occur, whereas larger pores disturb the stability of the scaffold (35). According to Dhandayuthapani et al. (17), and Chang et al. (36) optimal pore size depends on the nature of the cell (see Table 4). Pore interconnectivity must be sufficient to 8 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.

ensure that all cells are within 200 µm from the blood supply for mass transfer of oxygen and nutrients (17).

Table 4. Optimal pore size for cell infiltration according to Dhandayuthapani et al. (17) and Chang et al. (36)


Type of cell

Pore size (μm)

Red blood cell

5

Hepatocytes

20

Osteogenic cell

100-150

Adult mammalian skin cell

20-125

Smooth muscle cells

60-150

Endothelial cells

24

60-65

PDLLA

1200

44

12-16

55-60

+ biodegradable

- local reactions caused by acid degradation products - low mechanical properties

Elastomers PLA (74–76, 81, 82, 84, 111)

Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.

Tensile strength (MPa)

Degradation (month)

Tg (°C)

Advantages

Drawbacks

PGA (74, 76, 81, 82, 84, 111, 131)

6900

70

Several weeks

35-40

+ biodegradable

- low mechanical properties

PLGA (16, 84, 111)

1400-2800

41.4-55.2

1-12 (adjustable)

45-55

+ biodegradable + good degradation rate

- low mechanical properties

PCL (4, 16, 84–86, 111, 131)

0.21-0.34

20-34.5

>24

-72

+ good mechanical properties

- low degradation rate - low cell affinity

-60--6

+ good degradation + good processability

- low biostability - low cell affinity - carcinogenic - creep deformation

0.5

Several weeks

21


Tensile YM (MPa)

Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


Figure 7. Schematic of (a) end-on main chain, (b) side-on side chain, and (c) end-on side chain LCEs. Abbott et al. reported the effect of elastic stresses and defects of LC materials on the organization of certain lipids. They emphasized the connection between the mechanical properties of LC materials, LC ordering, and how this affected cell behavior (alignment) (134). Synthetically, the mechanical properties of LC materials can be tuned to match those of a range of cell types, and suitable approaches presented thus far include colloids in LC gels or LC gels derived from hydrogen-bonded molecular networks (135). Liquid crystal elastomers (LCEs) are soft materials that combine both the order of liquid crystals (orientational order and as a result anisotropic optical properties) (136) and the elasticity of elastomers (137, 138). Cross-linking allows LCEs to act as solids providing more support than gels or hydrogels. Nevertheless, the mesogenic groups within LCEs maintain considerable mobility because the network backbone formed by distant adjacent cross-links retains its rubbery state. Of the many potential applications for LCEs, sensors and actuators have been the most promising due to the LCEs ability to undergo significant shape change in response to a range of external stimuli (light, pH, temperature, among others) (139–144). The combination of biocompatibility, biodegradability and mechanical properties makes LCEs capable of satisfying most of the requirements of adaptable scaffolds, as we will discuss later. Also, the physical properties of LCEs can be tuned in a way to withstand mechanical tasks such as strain, stress, and impacts because they are soft, deformable (145), and can be functionalized with cell growth promoting moieties. They have also been found suitable as carriers in drug delivery applications. LCEs with their distinctive properties (146) 22 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


have been introduced as artificial muscles (liquid crystalline materials provide the greatest shape change with the least amount of energy) (147, 148), sensors (149, 150) and actuators (141, 151, 152), as tunable lasers (153), and as light-driven motors (154). LCEs are classified according to their structure. In a LCE, the mesogenic moieties may be attached into the main-chain (backbone), being part of the main chain, or be linked as a side (pendant) group on the main chain (see Figure 7). While LCEs can also show nematic or smectic order, the anisotropic arrangement (without an external stimuli) appears in form of a polydomain organization. Among the most promising properties of LCEs is the possibility to macroscopically orient the sample by mechanical stress, i.e., uniformly aligned the director and giving place to a monodomain organization within the structure. A stable monodomain arrangement can be reached by fixing the order by an extra cross-linking step as mentioned by Fleischmann et al. (146). Some of the factors that affect LCE behavior under stress (i.e., their Young’s modulus) are the ordering of the LC phase (e.g., nematic or smectic), the type of connectivity between LC molecules with respect to the polymer backbone (Figure 7), and the mono- or polydomain nature fixed during cross-linking. The LCEs’ anisotropic response to external stimuli (e.g., at the transition from the LC to the isotropic liquid phase) set them apart from conventional elastomer materials. Their soft and malleable nature allows them to conform to different shapes and surfaces (155, 156). The first nematic fixed monodomain elastomer was synthesized by Bergmann et al. (157). Recently, Yakacki et al. (158, 159) and Kim et al. (160) develop programmable monodomain LCEs. A potential advantage of LCEs is the often straightforward synthetic access to many structural variations and elastic properties from commercially available starting materials, which permits scale-up and high reproducibility. Several reports indicate that the contractile and expansion properties of these materials are associated with Young’s moduli (measured by tensile tests) between tens of kPa to several MPa; as such, LCEs can be regarded as artificial muscles (161–164) or used as biological actuators (146, 165). This range of elasticity perfectly matches those of several soft tissues in the human body. Considering both transverse Young’s moduli (obtained by indentation measurements) as well as Young’s moduli obtained by tensile measurements, the values range from 0.2 kPa for gray matter (darker tissue of brain and spinal cord) measured by indentation to ~480 and ~560 MPa for muscle and tendons, respectively. For LCEs to be considered as new materials for cell scaffolds, in addition to their optical and mechanical properties, they should also have well-defined porosity and surface properties to provide support for cell adherence, growth, and mass transport in and out of the scaffold under physiological conditions. Elastomer porosity should stimulate 3D cell–elastomer interactions, space for extracellular matrix (ECM) formation (6), and provide opportunities for linking molecular entities that allow binding of proteins, growth factors or proteins to enhance cellular adhesion. Considering all necessary requirements for LCEs as responsive cell scaffolds, we aimed at taking full advantage of the exceptional properties of the scaffolds’ liquid crystalline properties to guide cell attachment, proliferation, 23 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.

and alignment. Additional advantages result from built-in morphological cues (porous architecture) and their anisometric response to external stimuli. To do so, we prepared series of smectic and nematic LCEs that have proven to be non-cytotoxic to soft tissue cell lines. We have tested several selected cell lines (myoblasts, fibroblasts, and neuroblastomas) to create novel dynamic and spatial in vitro systems to simulate and study the development of new tissue as well as the complex cell-elastomer interplay.


Smectic-A (Sm-A) Liquid Crystalline Elastomers (LCEs) As a proof-of-concept, we reported the first synthesis, characterization and use of biocompatible, biodegradable, and porous LCEs (128) using a modular, convergent solvent-free synthesis (no impurities or toxic solvents), that allowed us to adjust porosity, degradation rate, and hydrophilic/hydrophobic balance. To prepare these new LCEs we combined three monomers, two lactone-based monomers and (D,L)-lactide, into a cross-linkable star block copolymer (SBC). Glycerol was chosen as central node due to its known nontoxicity and multifunctional reactive nature of the hydroxyl groups. ε-Caprolactone (ε-CL) and (D,L)-lactide are polymerized from this central node giving random chains fused by easily hydrolysable ester bonds known to produce biocompatible six-carbon fragments (highly important for biodegradation (166)). At a later stage, other polyols (building blocks of triglycerides) were chosen as central nodes offering access to series of three-, four- and six-arm SBCs (126, 167). Following a slightly modified method, involving ring opening polymerization (ROP), developed by Amsden et al. (126), the method modifications allowed for further chemical manipulation of the SBC via halogen atom substitution on the ε-caprolactone segments in two positions (α or γ) permitting a later attachment of LC pendant groups via alkyne-azide Huisgen’s cycloaddition reaction (“click” reaction) (128). Cholesterol groups were selected as LC pendants since cholesterol is known to be non-toxic and biocompatible. In fact, cholesterol is found in the organization of DNA (e.g., mitochondrial DNA-protein complexes attach to cholesterol-rich membrane segments) (168), the shell of certain shellfish and insects, and in cell walls of plants. Cholesterol is one of the first compounds found to be liquid crystalline; its liquid crystalline behavior was first described by Friedrich Reinitzer in 1888 (169, 170). Changing the position of the cholesterol groups in the substituted ε-caprolactone from the α- to the γ-position increases their mobility within the elastomer chain by reducing steric constrains (Figure 8). A bis-caprolactone was selected as a cross-linker to maintain identical, hydrolysable chemical bonds throughout the entire LCE as those forming the polymer backbone. LCEs prepared using this method can be casted or molded, and present an inherent porosity (spherical voids) due to the cross-linking process. The size and density of spherical voids is affected by position of the cholesterol pendants (α or γ), by the degree of cross-linking, and by the degree of cholesterol functionalization (ranging from 10 to 40%) with respect to the ε-caprolactone block. 24 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


Figure 8. (a) Chemical structure of 3-arm, 4-arm, and 6-arm initiators (central nodes), (b) synthesis pathway to star block copolymer-cholesterol liquid crystal (SBCα-CLC) and SBCγ-CLC, and (c) Crosslinking with bis-caprolactone (BCP) to obtain a 3-LCE-α or 3-LCE-γ. For better figure description see publication (167). Reproduced from reference (167). Copyright 2017, Macromol. Biosci., Wiley-VCH Verlag GmbH & Co. KGaA.

We have further adjusted the hydrophilic-hydrophobic balance (171) required for different cell types. This is easily achieved by introducing additional hydrophilic poly(ethylene oxide) (PEO) segments (Figure 9) instead of the 3-, 4- and 6-arm polyol central nodes. In contrast to the polyol nodes, introduction of PEO results in linear block-co-polymers (LBCs). Adjusting the hydrophobic-hydrophilic balance of the LCEs is important because certain cells 25 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


types prefer more hydrophilic surfaces than others. Myoblasts, for example, prefer more hydrophobic and neuroblastomas more hydrophilic LCE scaffolds (172).

Figure 9. (A) Cross-Linking of the polycaprolactone-based LC block-co-polymers using bis-caprolactone (CL = cross-linker) resulting in the formation of liquid crystal elastomers (LCEs), and (B) procedure for fabricating LCE foams with primary porosity (LCEFPP, Path 1) and LCE foams with both primary and secondary porosity (LCEFP+SP, Path 2). Reproduced from reference (172). Copyright 2016, ACS Macro. Lett.

26 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


Following the requisite of adjustable porosity, another type of porosity was accomplished using commercially available metal (Ni) foams as templates. This was performed during thermal cross-linking to create more porous LCEs with interconnected channels (tubular morphology) (172). These templates can be used in two ways, either fully immersed during cross-linking (tubular morphology), or quickly dipped into the pre-elastomer mixture prior to thermal cross-linking (see Figure 9). The latter produces a more complex architecture leading to LCE foams with primary and secondary porosity once the Ni template is etched away. These foams can be shaped in countless ways to obtain rolls, films, and folded constructs, among others.

Nematic Liquid Crystalline Elastomers (LCEs) De Gennes proposed the idea to think of nematic gels and LCEs as artificial muscles (147, 173). However there were no reports on the study of nematic LCEs as muscle cell scaffolds. To address this, we reported the first nematic LCEs as scaffolds for muscle cells and other tissues (174, 175). Nematic LCEs were prepared featuring a porous morphology using micro-emulsion photopolymerization with commercially available reactive mesogens (phenylbenzoate, PhBz) (176, 177). The reactive mesogens were confined within surfactant micelles, and the photopolymerization reaction captured the building blocks in a globular structure. Once photopolymerization was complete, the water–toluene solvent mixture was removed, and then the resulting LCEs scrupulously washed and rinsed to entirely eliminate the surfactant, which results in voids between the LCE globules. The resulting nematic LCE-PhBz exhibited a globular morphology with a porosity that proved suitable for seeding, growth, and proliferation of C2C12 myoblasts and other cell lines (see Figure 10). The overall LCE-PhBz porous morphology is suitable as 3D cell scaffold for several cell lines and allowed for satisfactory management of mass transport (i.e., nutrients, waste, and gases) even during longitudinal cell studies.

Mechanical and Thermal Properties The thermal properties for all LCEs were tested by differential scanning calorimetry (DSC) and thermal gravimetric analysis (TGA). All LCEs appeared amorphous lacking endothermic peaks indicative of melting. Predictably all Tg values were found to be considerably below physiological temperatures (128, 167), and LC phase formation and associated anisotropic mechanical properties were found from well below room to slightly above physiological temperatures (25 – 45 °C). 27 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


Figure 10. Chemical structure, phase transition temperatures of components, and ratio of monomer (M), crosslinker (L), and photoinitiator (PI; 2,2-dimethoxy-2-phenylacetophenone) used for the synthesis of the globular nematic LCE cell scaffolds. The cartoon depicts the synthesis (microemulsion photopolymerization) and the internal structure of the resulting fused globular LCEs. Reproduced from reference (174). Copyright 2015, ACS Appl. Mat. & Interfaces.

Substituting the 3-, 4- and 6-arm polyol central nodes considerably changes both the elasticity and the cell viability. The polyol central nodes can be seen as cross-linkers, where the 3-arm, and by analogy, the 6-arm LCEs have shown nearly identical elasticity. Contrary, the 4-arm SBCs lead to LCEs with higher Young’s moduli (~4 MPa). This can be related to previous theoretical and experimental studies where tetra-arm polymer hydrogel systems have extremely high homogeneous packing and suppressed heterogeneity, explaining the higher stiffness of the 4-arm SBC-based LCEs (LCE-4γ). Most of the studied LCEs have presently lower moduli (~2.0 – 4.0 MPa) than those of the selected tissues for study (~30 MPa for skin and ~480 MPa for muscle; see Table 3) (20), and some widely used biodegradable polymers. This is most likely due to the low molecular weight and nonlinear star-block structure (20, 167). However, the preliminary results of mechanical tests show an encouraging future for the use of LCEs as scaffolds. Cells were not only able to expand and proliferate, they aligned on and within the LCE scaffolds before any use of external stimuli. Figure 11 shows a comparative scale of the Young’s moduli of several soft tissues and of our LCEs measured by indentation and tensile tests. Figure 11 displays 28 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


that structural variety coupled with the choice of morphology allows our LCEs to cover a broad range of target tissue elasticity in both sets of measurements.

Figure 11. YM values obtained: (a) by indentation, and (b) by tensile measurements. Comparison of these values with measured YM values for specific tissues.

Recent synchrotron SAXS studies confirmed an increase in the degree of ordering of our LCE scaffolds (i.e., reorientation and alignment of the mesogenic units) upon applying mechanical stress. Further experiments to support this are now performed using immersion tensile stage measurements to study anisotropic changes of the LCE scaffolds immersed in media. The next step will then be to analyze the cellular response to the external mechanical stress applied when embedded in the LCE scaffold. In the case of LCE foams, there are several means to apply external stimuli, such as mechanical stress and mechanical compression as shown for the films. Other opportunities necessitate slight changes during preparation of the LCE foams such as adding magnetic nanoparticles (for magnetic stimuli), the addition of acrylate ends for an additional cross-linking step after mechanical induced stress, or pH-dependent moieties that could trigger a shape change.

3D Morphology The LCE scaffolds were prepared with three distinct morphologies: molded scaffolds or films, foams, and globular. Molded or film scaffolds were characterized by a morphology similar to that of a “Swiss Chesse”, foams (with two types of porosity: primary porous and primary as well as secondary porous), and globular (Figure 12). 29 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


Figure 12. SEM images of the internal porosity of: smectic LCE-γ (A) before, (B) and after 16 weeks of biodegradation in PBS, (C-D) of globular, porous microstructure of nematic LCE (E) LCEPP, (F) LCEPP, (G) Optical image of LCEFP+SP. Scheme of type de foams used to fabricate: (H) LCEPP , and (I) LCEFP+SP. Scanning electron microscopy (SEM) shows that pore size of the LCE films ranges from 10 to 30 µm in the dry state, and can potentially increase to about 20 to 60 µm once immersed in buffer or cell culture media (cells range in size from 5 to 150 µm in diameter, see Table 4) (17, 36). The second type of porosity achieved using commercially available metal (Ni) foams resulted in LCEs with interconnected channels or LCEs foams with primary and secondary porosity once the Ni template was etched away. The third morphology achieved using micro-emulsion photopolymerization with commercially available reactive mesogens featured spherical microparticles bonded together. These nematic LCEs exhibit a globular morphology (Figure 12 C and D). Thus far, these morphologies can be ranked with an increasing void volume as follows: the Swiss-cheese-type morphology ranges from 10 to 25%, the globular morphology from ~20 to 30%, the tubular morphology from 30 to 40%, and the LCE foams well above 70% as determined by Brunauer-Emmett-Teller adsorption isotherms (BET analysis). Since we can adjust this synthetically, this range of void volume can be tailored for specific tissue target microenvironments. For example, tubular morphologies can mimic microvessels in the brain. More open foam morphologies would be better for cell co-culturing of neural and glial cells or for engineering muscle tissue by co-culturing of myoblasts, fibroblasts and endothelial cells. The Swiss-cheese morphology is of interest for tougher implantable scaffolds. 30 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.

Mesomorphic Properties


LCEs are in general highly viscous and the preparation of thin-film polarized optical microscopy (POM) samples (sandwiched between untreated glass slides) proved often to be difficult. SAXD on the other hand allowed to make clear phase assignment. The presence of cholesterol groups in polymer side chains is known to promote cholesteric (chiral nematic) phases. However, all LCEs, except the nematic globular morphology, showed the presence of two scattering maxima with q1:q2 = 1:2 ([100] and [200]) values in the medium angle region denoting an ordered layer structure (Figure 13) with an interdigitated arrangement where all cholesterol molecules are packed side-by-side. This type of arrangement corresponds to an interdigitated smectic-A phase (SmA).

Figure 13. SAXD patterns of: A) LCE-γ, and B) LCE-α, C) model of the molecular arrangement of the pendant cholesteric LC groups matching the experimentally observed layer spacing. Figure 13C, reproduced from reference (128). Copyright 2015, Macromol. Biosci., Wiley-VCH Verlag GmbH & Co. KGaA. Cell-Elastomer Confocal Imaging Three cell lines were selected to test the capability of the LCE materials as cell scaffolds. Since LCEs have been thought of as artificial muscles, myoblasts (C2C12s) (178) were the first cell line to test; other cell lines were neuroblastomas (SH-Sy5Y) and primary dermal fibroblasts (hDF). Neuroblastomas (SH-SY5Y) derived from bone marrow differentiates into neural-like cells via retinoic acid, and primary human dermal fibroblasts are isolated from adult skin fibroblasts (hDF). These cell lines exemplify various soft tissues in the human body. LCEs films were obtained (representing a “Swiss-cheese” like morphology) from molding into ampoules, then carefully microtomed to obtain 60-, 100-, or 31 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


200-μm thin slices. The LCEs with tubular or foam morphology were used as prepared. Globular LCEs were briefly soaked in poly-D-lysine to promote cell adherence and temporarily mask the elastomer’s hydrophobic surface. Prior to cell culture studies, all LCEs were first sterilized by washing with 70% ethanol (or exposed to O2 plasma as an alternative). The LCE casted as thin films or foams were inserted in cell culture wells, and cells were seeded on and within the LCE scaffolds using the appropriate cell culture medium. After seeding, the cells were allowed to attach and grow for an initial period of 24 to 48 hrs inside an incubator (at 37 °C and 5% CO2). Proliferation (and potentially alignment) was monitored and periodically imaged by fluorescence confocal microscopy (FCM) for extended periods (days, weeks). All LCE cell-elastomer cultures were subjected to cell viability, proliferation, and cytotoxicity tests and assays (PrestoBlue cell viability, CyQuant cell proliferation, and Cyto Tox-Fluor cytotoxicity assay). All results from the combined assays and tests demonstrated that all LCE scaffolds clearly promoted cellular viability and proliferation without any inherent cytotoxicity (128, 167, 174). All LCE cell cultures (after staining with appropriate fluorescent dyes) were imaged over specific time intervals by FCM to analyze cell proliferation, density, and alignment. LCE-cell cultures showed remarkable features such as spontaneous cell alignment that largely depended on the morphology as well as the mechanical properties of a given LCE scaffold. For example, hDF cells grown on the 3-, 4-, and 6-arm LCEs showed highly anisotropic orientational behavior and proliferation on the 4- and 6-arm LCEs but not on the 3-arm LCEs (Figure 14). Fibroblasts are known to adhere better to substrates with contact angles between 60° and 80°, which are values matched only by the 4- and 6-arm LCEs (see Figure 14) (179, 180). Also the 4-arm LCEs usually have a higher Young’s moduli than the 3-arm LCEs (~4 MPa vs. ~2 MPa), with the 6-arm LCEs showing Young’s moduli somewhere in between (167). However, cells grown on the 6-arm LCEs also show extended cell nuclei that are characteristic of a cellular response to substrate strain (181) and/or topographical cues (182). The alignment of cell nuclei and its deviation from a spherical shape was analyzed using NIH’s ImageJ (183, 184).

Figure 14. Directionality analysis of primary human dermal fibroblast (hDF) cells grown on a) 3LCEα , b) 4LCEα , and c) 6LCEα elastomer films. The insets in each of the images show a photoimage and value from contact angle measurements. Reproduced from reference (167). Copyright 2017 Macromol. Biosci., Wiley-VCH Verlag GmbH & Co. KGaA. 32 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


The LCE foams shown in Figure 12 feature a 3-dimensional scaffold and while C2C12 cells commonly prefer more hydrophobic scaffold surfaces, they attached and proliferated on the inside walls of even the more hydrophilic LCE foam scaffolds with interconnected hollow channels. More importantly, cells within straight channel sections tended to spontaneously align parallel to the channel walls. This alignment tendency of the cells along channels, clearly indicated by elongated nuclei (184, 185), is also retained in curved channel sections (Figure 15). In contrast, in the center of the channel junctions, cells appeared randomly oriented, as indicated by round cell nuclei. C2C12 cells within this foam LCE scaffold seemed to thrive in the more open 3D microenvironment; they can freely interact with neighboring cells without randomly growing on top of one another (commonly observed in 2D cell scaffolds) or facing any spatial constraints.

Figure 15. Fluorescence confocal microscopy images of myoblast cells (C2C12) cultured in SBC-based LCEFP+SP using DMEM as a cell culture medium and stained with DAPI (for cell nuclei): (A) 2D images stacked in z direction and (B−D) 3D views from different angles. The image in (B) is plotted with the original fluorescence intensity value (original colors can be seen on publication). Reproduced from reference (172). Copyright 2016, ACS Macro. Lett.

LCE foams offer the possibility to engineer any free macroscopic form by molding the template into stripes, rolls, or other shapes, which allows for the construction of specific spatial cell environments to study not only cell-material but also cell-cell interactions, critical to emulating endogenous tissue. 33 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


The globular LCEs permitted cell adhesion directly onto the globular LCE surfaces and cells continued to grow and proliferate within the bulk of the globular elastomer matrix (Figure 16) that was co-stained to distinguish between cells and LCE matrix. It appears that the surface roughness and the porosity allowed the cells to attach better and permeate into the elastomer matrix. These nematic globular LCEs were also found to be excellent scaffolds for other cell lines such as hDF and SHSY5Y. Both types of SHSY5Y cells can be seen to have attached throughout the globular LCE scaffolds and stretching across several microspheres highlighting their expansion within the matrix and showing morphological phenotypes indicative of matured cells (186). Numerous neuritic extensions can be seen that are typical of these cell types (Figure 17).

Figure 16. Confocal microscopy images (x,y-, and x,z-plane) of C2C12 myoblasts 3, 5, and 7 days after seeding, costained with DAPI for cell nuclei and with rhodamine for the LCE scaffold. Reproduced from reference (174). Copyright 2015, ACS Appl. Mat. & Interfaces. 34 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


Figure 17. SEM images of: (A, ii–iv) C2C12 and (B, ii–iv) SHSY5Y after 7 and 5 days seeding, respectively. The cells (see arrows, for colored view see reference 169) can be seen as extending fibers directly attaching to the matrix for expansion and proliferation. Reproduced from reference (175). Copyright 2016, Frontiers, Loop (nature publishing group – npg) open-access Creative Commons License (CC-BY). Combining LCEs as scaffolds with 3D imaging techniques; fluorescence confocal microscopy and CARS (coherent anti-stokes Raman spectroscopy) will allow for real-time imaging and spectroscopic characterization of cells in 3D as they proliferate and differentiate under the various environmental/scaffold conditions. Additionally, the techniques will offer a solution to track multi-cellular interactions for extended periods of time. This in turn will help assess cell culture integrity by monitoring cell number, layering, longevity, and expansion into the scaffold to quantify cell extensions, developmental stages, and maturation markers. Degradation All LCEs based on ε-caprolactone-D,L-lactide start to fully in vitro biodegrade (under physiological pH) after a period of about 10 weeks (128). LCEs degrade at a faster rate when exposed to acidic or basic media (pH 3 or pH 11). The degradation process follows a bulk acid-catalyzed hydrolysis-long enough to permit cell alignment, and tissue formation. The degradation is autocatalytic, due to the formation of oligocarboxylic acids as the hydrolysis products. The LCEs first experience an increase in weight in the first week due to swelling (water absorption) into the bulk (Figure 18). It is known that water diffusion into caprolactone-lactide based elastomers decreases (at temperatures above Tg) as the elastomer cross-linking density increases (145). It is expected that the weight gain throughout the degradation studies will increase periodically due to the formation of degradation products within the matrix. This is turn will enhance the degradation and draw more water into the polymer matrix 35 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


(via osmosis) before the resulting degradation products finally leach out of the elastomer slab. Interestingly, LCE-α degraded faster than the non-LC elastomers, showing that the biodegradation rate is also affected by LC modification in the polymer backbone.

Figure 18. Biodegradation plots (time vs. % weight change) for: (A) LCE-α and (B) unmodified elastomer. Reproduced from reference (128). Copyright 2015, Macromol. Biosci., Wiley-VCH Verlag GmbH & Co. KGaA. During the degradation process the LCE scaffolds lose their integrity, some of its mechanical properties, and LC characteristics as cells continue to grow, mature, and proliferate, i.e., overtake the space once occupied by the LCE scaffold. Most of the LCE scaffold properties such as chemical composition, morphology, and mechanical properties affect cell fate. The scaffold modifications already made allow us to control chemical and morphological parameters to design responsive cell scaffolds that promote guided cell attachment, proliferation, and alignment, all the way to differentiation into specific cell lineages solely by morphological or mechanical cues native to the LCE scaffold.

Final Remarks and Chapter Conclusions LCE scaffolds containing a build-in anisotropic response to external stimuli together with the synthetic freedom to be able to adjust almost every chemical, mechanical and morphological parameters allow us to enhance our fundamental understanding of 3D cell microenvironments and cell response to scaffolds mimicking native tissues. Careful selection of elastic properties within modified LCEs will define the intended cell type to be grown to provide the most optimal proliferation rates. Looking forward into the future of LCEs as scaffolds there are several aspects that still need attention, such as of the temperature-dependent and 36 Ito et al.; Advances in Bioinspired and Biomedical Materials Volume 2 ACS Symposium Series; American Chemical Society: Washington, DC, 2017.


milieu-dependent mechanical properties of our LCE scaffolds. This includes measuring the Young’s moduli at physiological temperatures and conditions. Particular emphasis must be placed on conditions existing in vitro in cell culture media and in vitro in soft tissues (immersion biodegradability studies under stress), and in longitudinal tests at various stages of biodegradation, in parallel with cell culture. Regarding other morphologies, electro-spinning or melt-spinning are currently being tested and depend on the molecular structure of the LCE. In that case acrylate-based LCEs appear suitable and can undergo reactive electro-spinning with in situ photo-cross-linking by UV irradiation, not only adding to different size porosities but also interesting elastomer anisotropies. The success of a tissue-engineered material depends on many parameters, each of which must be carefully adjusted and controlled (i.e., elastic properties) to ensure cellular growth and optimal proliferation rates. Also, any 3D scaffold implanted at a defective or malfunctioning site must comprise the basic requirements of being biocompatible, occasionally biodegradable, appropriately porous, and be the location of cell attachment and proliferation. Overall, our 3D-LCE combine all those requirements together with the properties of LC materials providing a unique test platform that opens the path to improve current and develop new future cell therapies, as well as improved scaffolds for tissue regeneration. 3D LCE scaffold properties will impact molecular/cellular research and deliver new ways to understand and provide fundamental knowledge on how local microenvironments affect, for example, cell structure and protein expression. Our LCEs will help provide new insights into plastic/developmental changes occurring between certain types of cells, help us to determine which compounds cause these effects, and elucidating a mechanism to monitor these events. We envision incorporating other cell types and constituents into the scaffold including macrophages, endothelial cells, mast cells and microglia. There is sufficient pre-clinical evidence that stem cell therapies can result in major breakthroughs to cure or arrest the progression of many disease states and help prolong life (171). There are devastating degenerative diseases that cannot simply be cured by the use of therapeutic drugs. Stem cells have great potential to replace or repair damaged tissue, and can reverse degenerative diseases (diabetes, heart, lung, and neurological disorders) without the risk of side effects, as long as ways to implant these stem cells are created using scaffolds that mimic various affected tissues.

Dedication Je voudrais dédier ce chapitre à Patrick Prévôt, mon père – MP

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