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Mechanical mismatch. Stress−elongation curves of assorted (a) biological tissues and (b) polymeric gels, which demonstrate tissue's much stronger stiffening (Tables S1 and S2). Lines guide the reader, while data points represent literature data. (c) Stress− elongation responses of omental adipose tissue and silicone gel extracted from a commercial breast implant display a significant difference in strain-stiffening (β) despite a similarity in the Young's modulus (E 0 ). (d) An E 0 vs β map partitions polymeric gels (△) and biological tissues (○) as two distinct classes of materials. The β values are obtained by fitting stress−elongation curves with eq 1, whereas E 0 corresponds to the curve slope at λ → 1 (eq 2). The model is successful in fitting the entirety of gel elasticity, but only the elastic portion of tissue response before yielding. 17 Numbers at data points correspond to the stress−elongation curves in (a, b). Bottlebrush elastomers ( ■ ) mimic the stress−strain response of gels, 18 but are unable to reach the tissue territory.
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Softness and firmness are seemingly incompatible traits that synergize to create the unique soft-yet-firm tactility of living tissues pursued in soft robotics, wearable electronics, and plastic surgery. This dichotomy is particularly pronounced in tissues such as fat that are known to be both ultrasoft and ultrafirm. However, synthetically replicat...
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... tissues are distinct in possessing an oxymoronic mechanical property combination: they are compliant to the touch yet resistant to deformation, which imbues their characteristic feeling of firmness. 1,2 While initially they are very soft with Young's moduli ranging from E 0 = 10 3 −10 5 Pa, tissues rapidly stiffen by a factor of 10 2 −10 3 within a short interval of strain ( Figure 1a). 3−6 This strain-adaptive stiffening represents one of nature's key defense mechanisms that prevents accidental tissue rupture and serves as a benchmark for various industrial 7,8 and biomedical applications. ...
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... This strain-adaptive stiffening represents one of nature's key defense mechanisms that prevents accidental tissue rupture and serves as a benchmark for various industrial 7,8 and biomedical applications. 9−12 Tissue softness is routinely replicated with polymer gels, 13 but gels are limited in their ability to copy nature's strain-stiffening capabilities ( Figure 1b) as accentuated by the sharply diverging deformation responses of adipose tissue and a silicone gel utilized in breast implants ( Figure 1c). This mechanical mismatch is further exacerbated by both spontaneous 14 and induced 15 solvent migration leading to inadequate performance of engineered devices and unforeseen health risks. ...
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... This strain-adaptive stiffening represents one of nature's key defense mechanisms that prevents accidental tissue rupture and serves as a benchmark for various industrial 7,8 and biomedical applications. 9−12 Tissue softness is routinely replicated with polymer gels, 13 but gels are limited in their ability to copy nature's strain-stiffening capabilities ( Figure 1b) as accentuated by the sharply diverging deformation responses of adipose tissue and a silicone gel utilized in breast implants ( Figure 1c). This mechanical mismatch is further exacerbated by both spontaneous 14 and induced 15 solvent migration leading to inadequate performance of engineered devices and unforeseen health risks. ...
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... guide the materials design toward soft tissue firmness, we introduce an equation of state relating true stress σ true with sample uniaxial elongation ratio λ = L/L 0 from its initial size L 0 to deformed size L as which describes the nonlinear elastic response of polymer networks ( Figure 1a,b) as a function of two molecular parameters: structural modulus E and strain-stiffening parameter β. 19 The modulus is controlled by the density (ρ s ) and conformation of stress-supporting strands as E ≅ k B Tρ s ⟨R in 2 ⟩/ (b K R max ), where b K , ⟨R in 2 ⟩, and R max are a strand's Kuhn length, mean square end-to-end distance, and contour length. The strain-stiffening curvature (Figure 1c) is determined by potential extensibility of network strands from their initial mean-square end-to-end distance ⟨R in 2 ⟩ to the corresponding contour length of a fully extended strand as β = ⟨R in 2 ⟩/R max 2 such that 0 < β < 1. ...
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... guide the materials design toward soft tissue firmness, we introduce an equation of state relating true stress σ true with sample uniaxial elongation ratio λ = L/L 0 from its initial size L 0 to deformed size L as which describes the nonlinear elastic response of polymer networks ( Figure 1a,b) as a function of two molecular parameters: structural modulus E and strain-stiffening parameter β. 19 The modulus is controlled by the density (ρ s ) and conformation of stress-supporting strands as E ≅ k B Tρ s ⟨R in 2 ⟩/ (b K R max ), where b K , ⟨R in 2 ⟩, and R max are a strand's Kuhn length, mean square end-to-end distance, and contour length. The strain-stiffening curvature (Figure 1c) is determined by potential extensibility of network strands from their initial mean-square end-to-end distance ⟨R in 2 ⟩ to the corresponding contour length of a fully extended strand as β = ⟨R in 2 ⟩/R max 2 such that 0 < β < 1. For polymer networks with nonlinear responses, the Young's modulus E 0 depends not only on network strands density (E ∼ ρ s ), but also on their initial ...
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... November 25, 2019 Published: January 22,2020 Simply, this elastic model provides two parameters observable in stress−elongationplots (Figure 1c): the initial slope or softness (E 0 ) followed by a curvature or firmness (β), which characterizes resistivity of material to deformation. Respectively, mapping [E 0 , β] allows partitioning gels and tissues into two distinct materials classes with similar E 0 yet vastly different β (β gel ≪ β tissue ) ( Figure 1d). ...
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... November 25, 2019 Published: January 22,2020 Simply, this elastic model provides two parameters observable in stress−elongationplots (Figure 1c): the initial slope or softness (E 0 ) followed by a curvature or firmness (β), which characterizes resistivity of material to deformation. Respectively, mapping [E 0 , β] allows partitioning gels and tissues into two distinct materials classes with similar E 0 yet vastly different β (β gel ≪ β tissue ) ( Figure 1d). The limited firmness of linearchain polymeric gels (β gel = β dry ⟨R in 2 ⟩ gel /⟨R in 2 ⟩ dry = β dry α 2/3 < 0.2) originates both from weak strand extension in as-prepared networks (β dry ≅ 0.01) and an upper bound on their swelling ratio (α < 100). ...
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... For a deeper discussion on the origins and validity of this elastic model, we encourage the reader to pursue prior literature. 19 Various molecular and macroscopic constructs have endeavored to bridge the strain-stiffening divide and replicate tissue firmness (0.7 < β < 1) in Figure 1b, 17,18,21−26 but as of now, most attempts have fallen short of β > 0.4, including our earlier studies utilizing solvent-free bottlebrush elastomers (filled squares ■ in Figure 1d). 18 In this regard, self-assembled networks of linear−bottlebrush−linear (LBL) block copolymers have proven to be a resourceful scaffold given the hierarchical integration of molecular and particulate motifs within each network strand (Figure 2a). ...
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... For a deeper discussion on the origins and validity of this elastic model, we encourage the reader to pursue prior literature. 19 Various molecular and macroscopic constructs have endeavored to bridge the strain-stiffening divide and replicate tissue firmness (0.7 < β < 1) in Figure 1b, 17,18,21−26 but as of now, most attempts have fallen short of β > 0.4, including our earlier studies utilizing solvent-free bottlebrush elastomers (filled squares ■ in Figure 1d). 18 In this regard, self-assembled networks of linear−bottlebrush−linear (LBL) block copolymers have proven to be a resourceful scaffold given the hierarchical integration of molecular and particulate motifs within each network strand (Figure 2a). ...
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... To validate this concept, we synthesized two groups of LBL triblock copolymers with a polydimethylsiloxane (PDMS) bottlebrush block (B-block) and two poly(methyl methacrylate) (PMMA) linear end-blocks (L-blocks). These groups differ by the degree of polymerization (DP) of PDMS side chains, n sc = 14 (Group 1) and n sc = 70 (Group 2) and contain several series of LBL triblocks with different DP's of bottlebrush backbone (n bb = 100−1100) and PMMA Lblock (n L = 50−1300) (Table S3). Figure 3a compares representative stress−elongation curves from each group to demonstrate the effects of n L and n sc on the Young's modulus and strain-stiffening ( Figure S1 shows a complete set of deformation curves). An [E 0 , β] map reveals that all Group 1 plastomers coalesce (green, Figure 3b) to successfully cross the gel-tissue divide, yet skirt many essential tissues such as muscle (β = 0.7), skin (β = 0.8), and fat (β = 0.9) located in the bottom-right corner of the tissue territory. ...
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... 4a exemplifies stress−elongation curves of assorted plastomers that reveal agreeable mechanical responses with brain, fetal membrane, and spinal cord tissues. These solvent-free materials also successfully replicate adipose tissues ( Figure 4b) and serve as a superior solution to commercially available silicone gel-based products (Figure 1c) that leach into the body. 15,16 Additionally, biological tissues exhibit significant variation in mechanical response depending on bodily location, age, strain rate, and deformation direction with respect to tissue texture. ...
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... LBL platform also exhibits adequate biocompatibility as demonstrated by the adhesion and proliferation of human normal mammary epithelial and adipose-derived mesenchymal stem cells (MSCs) cultured onto a 300−1/70 surface ( Figure 4c). Monitoring the cultured cells by fluorescence microscopy over the course of a week reveals plastomers as adequate substrates for both cell's viability and proliferation (Figure Figure 1d where Group 1 plastomers (green) enable gel-tissue bridging, while Group 2 plastomers (black) successfully penetrate into the tissue territory (Tables S1−S3). Dashed lines are used to guide the reader and not indicative of theoretical correlation. ...
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... lines are used to guide the reader and not indicative of theoretical correlation. Group 2 plastomers with shorter backbones (n bb = 100) shift toward higher E 0 (black •), due to a starlike strand conformation as n bb ≈ n sc ( Figure S1). (c) Correlation between mechanical properties (E 0 and β) and molecular parameters (n L , n bb , ϕ L ) demonstrate good agreement with theoretical analysis summarized in eq S10, where ϕ=n g /(n g +n sc ). ...
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... each case, tests were conducted in triplicate to ensure accuracy of the data. All stress−elongationcurves show dependence of the true stress σ true on the elongation ratio λ in accordance with eq 1 at small and intermediate deformation range but switch to a linear scaling with λ at the later stages of deformation ( Figure S1 and Table S3). The elongation ratio λ for uniaxial network deformation is defined as the ratio of the sample's instantaneous size L to its initial size L 0 , λ = L/L 0 . ...
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... tissues are distinct in possessing an oxymoronic mechanical property combination: they are compliant to the touch yet resistant to deformation, which imbues their characteristic feeling of firmness. 1,2 While initially they are very soft with Young's moduli ranging from E 0 = 10 3 −10 5 Pa, tissues rapidly stiffen by a factor of 10 2 −10 3 within a short interval of strain ( Figure 1a). 3−6 This strain-adaptive stiffening represents one of nature's key defense mechanisms that prevents accidental tissue rupture and serves as a benchmark for various industrial 7,8 and biomedical applications. ...
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... This strain-adaptive stiffening represents one of nature's key defense mechanisms that prevents accidental tissue rupture and serves as a benchmark for various industrial 7,8 and biomedical applications. 9−12 Tissue softness is routinely replicated with polymer gels, 13 but gels are limited in their ability to copy nature's strain-stiffening capabilities ( Figure 1b) as accentuated by the sharply diverging deformation responses of adipose tissue and a silicone gel utilized in breast implants ( Figure 1c). This mechanical mismatch is further exacerbated by both spontaneous 14 and induced 15 solvent migration leading to inadequate performance of engineered devices and unforeseen health risks. ...
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... This strain-adaptive stiffening represents one of nature's key defense mechanisms that prevents accidental tissue rupture and serves as a benchmark for various industrial 7,8 and biomedical applications. 9−12 Tissue softness is routinely replicated with polymer gels, 13 but gels are limited in their ability to copy nature's strain-stiffening capabilities ( Figure 1b) as accentuated by the sharply diverging deformation responses of adipose tissue and a silicone gel utilized in breast implants ( Figure 1c). This mechanical mismatch is further exacerbated by both spontaneous 14 and induced 15 solvent migration leading to inadequate performance of engineered devices and unforeseen health risks. ...
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... guide the materials design toward soft tissue firmness, we introduce an equation of state relating true stress σ true with sample uniaxial elongation ratio λ = L/L 0 from its initial size L 0 to deformed size L as which describes the nonlinear elastic response of polymer networks ( Figure 1a,b) as a function of two molecular parameters: structural modulus E and strain-stiffening parameter β. 19 The modulus is controlled by the density (ρ s ) and conformation of stress-supporting strands as E ≅ k B Tρ s ⟨R in 2 ⟩/ (b K R max ), where b K , ⟨R in 2 ⟩, and R max are a strand's Kuhn length, mean square end-to-end distance, and contour length. The strain-stiffening curvature (Figure 1c) is determined by potential extensibility of network strands from their initial mean-square end-to-end distance ⟨R in 2 ⟩ to the corresponding contour length of a fully extended strand as β = ⟨R in 2 ⟩/R max 2 such that 0 < β < 1. ...
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... guide the materials design toward soft tissue firmness, we introduce an equation of state relating true stress σ true with sample uniaxial elongation ratio λ = L/L 0 from its initial size L 0 to deformed size L as which describes the nonlinear elastic response of polymer networks ( Figure 1a,b) as a function of two molecular parameters: structural modulus E and strain-stiffening parameter β. 19 The modulus is controlled by the density (ρ s ) and conformation of stress-supporting strands as E ≅ k B Tρ s ⟨R in 2 ⟩/ (b K R max ), where b K , ⟨R in 2 ⟩, and R max are a strand's Kuhn length, mean square end-to-end distance, and contour length. The strain-stiffening curvature (Figure 1c) is determined by potential extensibility of network strands from their initial mean-square end-to-end distance ⟨R in 2 ⟩ to the corresponding contour length of a fully extended strand as β = ⟨R in 2 ⟩/R max 2 such that 0 < β < 1. For polymer networks with nonlinear responses, the Young's modulus E 0 depends not only on network strands density (E ∼ ρ s ), but also on their initial ...
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... this elastic model provides two parameters observable in stress−elongationplots (Figure 1c): the initial slope or softness (E 0 ) followed by a curvature or firmness (β), which characterizes resistivity of material to deformation. Respectively, mapping [E 0 , β] allows partitioning gels and tissues into two distinct materials classes with similar E 0 yet vastly different β (β gel ≪ β tissue ) ( Figure 1d). ...
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... this elastic model provides two parameters observable in stress−elongationplots (Figure 1c): the initial slope or softness (E 0 ) followed by a curvature or firmness (β), which characterizes resistivity of material to deformation. Respectively, mapping [E 0 , β] allows partitioning gels and tissues into two distinct materials classes with similar E 0 yet vastly different β (β gel ≪ β tissue ) ( Figure 1d). The limited firmness of linearchain polymeric gels (β gel = β dry ⟨R in 2 ⟩ gel /⟨R in 2 ⟩ dry = β dry α 2/3 < 0.2) originates both from weak strand extension in as-prepared networks (β dry ≅ 0.01) and an upper bound on their swelling ratio (α < 100). ...
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... For a deeper discussion on the origins and validity of this elastic model, we encourage the reader to pursue prior literature. 19 Various molecular and macroscopic constructs have endeavored to bridge the strain-stiffening divide and replicate tissue firmness (0.7 < β < 1) in Figure 1b, 17,18,21−26 but as of now, most attempts have fallen short of β > 0.4, including our earlier studies utilizing solvent-free bottlebrush elastomers (filled squares ■ in Figure 1d). 18 In this regard, self-assembled networks of linear−bottlebrush−linear (LBL) block copolymers have proven to be a resourceful scaffold given the hierarchical integration of molecular and particulate motifs within each network strand (Figure 2a). ...
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... For a deeper discussion on the origins and validity of this elastic model, we encourage the reader to pursue prior literature. 19 Various molecular and macroscopic constructs have endeavored to bridge the strain-stiffening divide and replicate tissue firmness (0.7 < β < 1) in Figure 1b, 17,18,21−26 but as of now, most attempts have fallen short of β > 0.4, including our earlier studies utilizing solvent-free bottlebrush elastomers (filled squares ■ in Figure 1d). 18 In this regard, self-assembled networks of linear−bottlebrush−linear (LBL) block copolymers have proven to be a resourceful scaffold given the hierarchical integration of molecular and particulate motifs within each network strand (Figure 2a). ...
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... To validate this concept, we synthesized two groups of LBL triblock copolymers with a polydimethylsiloxane (PDMS) bottlebrush block (B-block) and two poly(methyl methacrylate) (PMMA) linear end-blocks (L-blocks). These groups differ by the degree of polymerization (DP) of PDMS side chains, n sc = 14 (Group 1) and n sc = 70 (Group 2) and contain several series of LBL triblocks with different DP's of bottlebrush backbone (n bb = 100−1100) and PMMA Lblock (n L = 50−1300) (Table S3). Figure 3a compares representative stress−elongation curves from each group to demonstrate the effects of n L and n sc on the Young's modulus and strain-stiffening ( Figure S1 shows a complete set of deformation curves). An [E 0 , β] map reveals that all Group 1 plastomers coalesce (green, Figure 3b) to successfully cross the gel-tissue divide, yet skirt many essential tissues such as muscle (β = 0.7), skin (β = 0.8), and fat (β = 0.9) located in the bottom-right corner of the tissue territory. ...
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... 4a exemplifies stress−elongation curves of assorted plastomers that reveal agreeable mechanical responses with brain, fetal membrane, and spinal cord tissues. These solvent-free materials also successfully replicate adipose tissues ( Figure 4b) and serve as a superior solution to commercially available silicone gel-based products (Figure 1c) that leach into the body. 15,16 Additionally, biological tissues exhibit significant variation in mechanical response depending on bodily location, age, strain rate, and deformation direction with respect to tissue texture. ...
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... LBL platform also exhibits adequate biocompatibility as demonstrated by the adhesion and proliferation of human normal mammary epithelial and adipose-derived mesenchymal stem cells (MSCs) cultured onto a 300−1/70 surface ( Figure 4c). Monitoring the cultured cells by fluorescence microscopy over the course of a week reveals plastomers as adequate substrates for both cell's viability and proliferation (Figure Figure 1d where Group 1 plastomers (green) enable gel-tissue bridging, while Group 2 plastomers (black) successfully penetrate into the tissue territory (Tables S1−S3). Dashed lines are used to guide the reader and not indicative of theoretical correlation. ...
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... lines are used to guide the reader and not indicative of theoretical correlation. Group 2 plastomers with shorter backbones (n bb = 100) shift toward higher E 0 (black •), due to a starlike strand conformation as n bb ≈ n sc ( Figure S1). (c) Correlation between mechanical properties (E 0 and β) and molecular parameters (n L , n bb , ϕ L ) demonstrate good agreement with theoretical analysis summarized in eq S10, where ϕ=n g /(n g +n sc ). ...
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... each case, tests were conducted in triplicate to ensure accuracy of the data. All stress−elongationcurves show dependence of the true stress σ true on the elongation ratio λ in accordance with eq 1 at small and intermediate deformation range but switch to a linear scaling with λ at the later stages of deformation ( Figure S1 and Table S3). The elongation ratio λ for uniaxial network deformation is defined as the ratio of the sample's instantaneous size L to its initial size L 0 , λ = L/L 0 . ...
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... Crosslinked polymer networks are important in diverse applications that range from membranes for selective transport to biological scaffolds to the matrix used in 3D-printing. [1][2][3][4] There are numerous strategies to control the properties of polymer networks; simple strategies include varying the crosslinking density or the polymer strand length, and more sophisticated strategies include tuning the architecture of the polymers that comprise the network [5][6][7] , creating interpenetrating double networks 8,9 , or creating microphase separation in the network that is locked in place by the crosslinking process 10 . Networks with microphase separation between rubbery and glassy polymers are a promising class of materials that are similar to block copolymers in that they can potentially create co-continuous structures with complimentary properties 10,11 . ...
... The domain spacing in real space units, calculated as d = 2π/q * , gives d = [7.8σ , 10.6σ , 15.8σ ] for N = [5,10,20], respectively. The scaling dependence of the domain spacing on molecular weight is fit to be approximately 0.51 ( Figure 1b, inset), which is consistent with the prediction of d ∼ N 0.5 by de Gennes for a polymer melt blend constrained by cross-linkers 43 . ...
... The high-T g phase data are fitted to the Arrhenius function (solid) and the low-T g phase data are fitted to the WLF equation (dashed). The three domain spacing sizes are indicated by the corresponding molecular weight of the chains N = (5,10,20) in blue, orange and green, respectively. ...
Inhomogeneous crosslinked polymers are a powerful platform for materials design, because they can be synthesized from materials that provide complimentary properties to the resulting gel. For example, a membrane with...
... However, it was discovered that special types of bottlebrush copolymers can closely mimic the peculiar mechanical properties of soft tissue [493]. To bridge the gap between mechanical properties of ordinary hydrogels and soft tissues, an ATRP strategy was used to make a linear-bottlebrush-linear (L-BB-L) architecture [494]. The middle PDMS bottlebrush segment was made by an ATRP strategy; then end chains were modified to serve as initiation sites for ATRP. ...
... The obtained PDMS macroinitiator with dual initiation sites underwent another ATRP to grow PMMA chains from the bottlebrush polymer, resulting in the L-BB-L architecture. These solvent-free elastomers were observed to closely mimic the strain-stress behavior of different soft tissues like brain and adipose [494]. ...
The continuing wave of technological breakthroughs and advances is critical for engineering well-defined materials, particularly biomaterials, with tailored microstructure and properties. Over the last few decades, controlled radical polymerization (CRP) has become a very promising option for the synthesis of precise polymeric materials with an unprecedented degree of control over molecular architecture. Atom Transfer Radical Polymerization (ATRP), one of the most robust and efficient CRPs, has been at the forefront of the synthesis of well-defined polymers with controlled/predetermined molecular weights, polydispersity, topology, composition, and site-specific functionality. ATRP has been leveraged to prepare a wide range of polymers with properties tailored for a number of biomedical applications. Furthermore, ATRP can also be utilized to introduce stimuli-responsive properties into the chemical structure of polymers. Moreover, the degradation behavior of ATRP-derived polymers can be tailored by incorporating chemical bonds susceptible to hydrolysis or proteolysis. This strategy allows the design of degradable polymers for in vivo applications. This review summarizes the recent advances in ATRP for the design of functional materials and techniques implemented to advance the biomedical field, such as surface modification and functionalization. Additionally, the latest applications and progress of ATRP-derived materials in various biomedical arenas such as drug delivery, tissue engineering, bioimaging, and biosensing are reported. Lastly, the current limitations and future perspectives of ATRP-derived biomaterials are carefully discussed to support further improvement of their properties and performance for translatability into the clinic. Moving forward, there is a need for further development of ATRP to align with green chemistry principles. This entails exploring the use of renewable monomers, environmentally friendly and nontoxic solvents, as well as metal-free and biocompatible catalysts. Additionally, researchers should thoroughly investigate the bioactivity, biodegradation behavior, and in vivo fate of ATRP-derived polymers and polymer conjugates before considering their translation into clinical applications.
... Simultaneously, BB elastomers reveal significant stiffening during deformation. This combination of initial softness and intense strain-stiffening is vital for designing biomimetic materials capable of replicating the mechanical properties of soft living tissues [24][25][26]. By introducing crystallizable side chains into such systems, another tuning mechanism becomes available, rendering these materials temperature-sensitive [27]. ...
Bottlebrush (BB) elastomers with water-soluble side chains and tissue-mimetic mechanical properties are promising for biomedical applications like tissue implants and drug depots. This work investigates the microstructure and phase transitions of BB elastomers with crystallizable polyethylene oxide (PEO) side chains by real-time synchrotron X-ray scattering. In the melt, the elastomers exhibit the characteristic BB peak corresponding to the backbone-to-backbone correlation. Upon crystallization of the side chains, the intensity of the peak decays linearly with crystallinity, and eventually vanishes due to BB packing disordering within intercrystalline amorphous gaps. This behavior of the bottlebrush peak differs from an earlier study of BBs with poly(ε-caprolactone) side chains, explained by stronger backbone confinement in the case of PEO, a high-crystallinity polymer. Microstructural models based on 1D SAXS correlation function analysis suggest crystalline lamellae of PEO side chains separated by amorphous gaps of monolayer-like BB backbones.
... Another approach uses linear-bottlebrush-linear (LBBL) triblock copolymers to create a physically-linked bottlebrush network, i.e. thermoplastic elastomer, which does not require a post-treatment crosslinking step. 5,[30][31][32] A PDMS-based LBBL thermoplastic elastomer operating at room temperature DIW printing in the way that the material is solvated with a volatile solvent that evaporates as the extruded material exits the nozzle, leading to solidification which provides enough mechanical strength for the printed feature. 29 Although the solvent facilitates a sharp solid-to-liquid transition at lower stress values, its use presents challenges. ...
Polymer networks containing bottlebrush chains are emerging materials with exceptionally soft and highly tunable mechanical properties. However, such materials have not been extensively implemented in functional processing techniques such as three-dimensional (3D) printing. Here, we introduce a new design of soft and solvent-free polydimethylsiloxane (PDMS)-based thermoplastic elastomer which contains dangling and space-filling bottlebrush chains, featuring a yield stress and a rapid recovery after stress removal; both required for high spatial fidelity 3D printing. The developed material is composed of two copolymers; the main building block is a diblock copolymer with linear polystyrene (PS) block and bottlebrush PDMS block (PS-b-bbPDMS) while the second component is PS-b-PDMS-b-PS triblock, self-assembling to a physical network. This design provides independent tunability of each structural parameter on the molecular level, hence, macroscopic control of the materials' mechanical properties. Multiple self-supportive 3D structures with spanning elements are 3D printed at elevated temperatures using a developed material with a low shear modulus of G′ = 3.3 kPa containing 3 : 1 molar ratio of diblock to triblock copolymers without the need for volatile solvent, or post-treatment. This 3D printing compatible design opens new opportunities to utilize the distinctive mechanical properties of bottlebrush materials for applications such as soft tissue scaffolds, sensors, actuators, and soft robots.
... Ultrathin polymer sheets or nanomeshes are nice examples that can effectively reduce mechanical constraints and increase conformability to arbitrary surfaces [25][26][27] , but the ultrathin structures with in-plane patterns require complex device fabrication process. Alternatively, bottlebrush elastomers (BBEs) are a class of intrinsically stretchable materials that can achieve ultra-low Young's modulus without solvents, because of their highly branched architecture consisting of polymeric side chains attached to a polymer backbone, leading to reduced entanglements in comparison to linear analogs [28][29][30][31][32][33] . For example, polydimethylsiloxane (PDMS)-based BBEs can achieve a Young's modulus of <1 kPa 28,34 , comparatively lower than that of commercial PDMS linear elastomers such as the Sylgard 184 with a Young's modulus of 100 kPa to 3 MPa 35,36 (Fig. 1b). ...
Understanding biological systems and mimicking their functions require electronic tools that can interact with biological tissues with matched softness. These tools involve biointerfacing materials that should concurrently match the softness of biological tissue and exhibit suitable electrical conductivities for recording and reading bioelectronic signals. However, commonly employed intrinsically soft and stretchable materials usually contain solvents that limit stability for long-term use or possess low electronic conductivity. To date, an ultrasoft (i.e., Young’s modulus <30 kPa), conductive, and solvent-free elastomer does not exist. Additionally, integrating such ultrasoft and conductive materials into electronic devices is poorly explored. This article reports a solvent-free, ultrasoft and conductive PDMS bottlebrush elastomer (BBE) composite with single-wall carbon nanotubes (SWCNTs) as conductive fillers. The conductive SWCNT/BBE with a filler concentration of 0.4 − 0.6 wt% reveals an ultralow Young’s modulus (<11 kPa) and satisfactory conductivity (>2 S/m) as well as adhesion property. Furthermore, we fabricate ultrasoft electronics based on laser cutting and 3D printing of conductive and non-conductive BBEs and demonstrate their potential applications in wearable sensing, soft robotics, and electrophysiological recording.
... Sheiko and co-workers reported A-B-A triblock copolymers with linear polymer A blocks and a graft copolymer B block; these structures formed very soft thermoplastic elastomers through microphase separation of their A blocks into spherical morphologies bridged by the B segment. 33 Bates and coworkers reported statistical graft copolymers featuring polydimethylsiloxane (PDMS) and smaller, oligo(ethylene oxide) (OEO) side chains; intramolecular phase separation of the PDMS and OEO segments yielded very soft, shearthinning materials with unbridged spherical morphologies. 39 In a different vein, Grubbs and co-workers explored the impact of small-molecule monomers as "diluents" in the formation of diblock bottlebrush copolymers, revealing how tapered and gradient side chains can impact diblock bottlebrush assembly. ...
Graft copolymers offer a versatile platform for the design of self-assembling materials; however, simple strategies for precisely and independently controlling the thermomechanical and morphological properties of graft copolymers remain elusive. Here, using a library of 92 polynorbornene-graft-polydimethylsiloxane (PDMS) copolymers, we discover a versatile backbone-pendant sequence-control strategy that addresses this challenge. Small structural variations of pendant groups, e.g., cyclohexyl versus n-hexyl, of small-molecule comonomers have dramatic impacts on order-to-disorder transitions, glass transitions, mechanical properties, and morphologies of statistical and block silicone-based graft copolymers, providing an exceptionally broad palette of designable materials properties. For example, statistical graft copolymers with high PDMS volume fractions yielded unbridged body-centered cubic morphologies that behaved as soft plastic crystals. By contrast, lamellae-forming graft copolymers provided robust, yet reprocessable silicone thermoplastics (TPs) with transition temperatures spanning over 160 °C and elastic moduli as high as 150 MPa despite being both unentangled and un-cross-linked. Altogether, this study reveals a new pendant-group-mediated self-assembly strategy that simplifies graft copolymer synthesis and enables access to a diverse family of silicone-based materials, setting the stage for the broader development of self-assembling materials with tailored performance specifications.
... In most cases, they are homogeneous materials, and their bulk properties are designed considering the final applications. Hydrogels are multipurpose soft matters that are used as scaffolds for tissue engineering (Lee and Mooney, 2001;Keith et al., 2020), vehicles for drug delivery (Qiu and Park, 2001;GhavamiNejad et al., 2016b), actuators for optics and fluidics (Dong et al., 2006;Vatankhah-Varnoosfaderani et al., 2017), model extracellular matrices for biological studies (Discher et al., 2009;Dutta et al., 2019), and chemical engineering applications, because of their outstanding biological and physical chemistry features such as biocompatibility, transparency, impact resistance, water absorbance, and separation capability (Annaka et al., 2003;Isik, 2003;Haraguchi and Takada, 2005). However, their mechanical behavior often severely limits the scope of hydrogel applications (Calvert, 2009). ...
Hybrid hydrogels based on n-isopropylacrylamide, zwitterionic comonomer, and graphene oxide were synthesized to study their physical and mechanical properties. The compositional variation largely influenced the swelling characteristics of the hybrid hydrogels compared to mechanical properties, i.e., elongation and compression. Additionally, Rheometric swelling measurements on the swollen hydrogels were performed until they reached equilibrium showed a very low phase angle δ indicating strong covalent network, which intrun increases with increasing content of zwitterions and GO. Swelling kinetics were studied and found to follow Fickian dynamics, albeit zwitterion-containing gels showed a peculiar 2-step swelling pattern. Interestingly, differences in the swelling mechanism are also clear for the hydrogels with 2D GO (Graphene oxide) nano-fillers from its 1D nano-filler CNTs (Carbon nanotubes). . In elongation, the samples break in a brittle fashion at Hencky strains εmax around 0.4-0.65 with the maximum stress being observed for samples with high Zw-content and 0.2% GO, which can be explained by the stress-rising properties of sharp edges of GO. In contrast, the data in compression profits from higher GO-contents as crack growth is less important in this deformation mode. This work will contribute to future composite gel applications.
... Architectural parameters and mechanical properties of A-g-B brush copolymers. (1) Grafting density of side chains on the backbone with BA spacer. (2) Number average degree polymerization of brush backbone between glassy block side chains that physical crosslink. ...
... Sample 090320_4: . S11.1 H-NMR of purified poly[nBA-ran-MMA-g-(PDMS/PMMA)] brush copolymer series =1923, =4 synthesized by CRP (400 MHz, CDCl3): 3.89 ((C=O)-O-CH2-, PDMS brush, m, 2H), 3.62 ((C=O)-O-CH3-and -O-CH2-CH2-, PMMA and PEG respectively, m, 7H), 0.09 (-(Si(CH3)2-O)n-, PDMS macromonomer and brush mixture, s, 68.4H). 19 Fig. S12. 1 H-NMR of purified poly[nBA-ran-MMA-g-(PDMS/PMMA)] brush copolymer series =1959, =8 synthesized by CRP (400 MHz, CDCl3): 3.89 ((C=O)-O-CH2-, PDMS brush, m, 2H), 3.62 ((C=O)-O-CH3-and -O-CH2-CH2-, PMMA and PEG respectively, m, 7H), 0.09 (-(Si(CH3)2-O)n-, PDMS macromonomer and brush mixture, s, 68.4H). ...
... S17.1 H-NMR of purified poly[nBA-ran-MMA-g-(PDMS/PS)]. (400 MHz, CDCl3): 6.56 (CH-CH=CH-C-, PS side-chain, d, 120H), 4.07 ((C=O)-O-CH2-, PBA spacer, d, 2H), 0.09 (-(Si(CH3)2-O)n-, PDMS macromonomer and brush mixture, s, 68.4H). ...
Polymeric networks are commonly used for various biomedical applications, from reconstructive surgery to wearable electronics. Some materials may be soft, firm, strong, or damping however, implementing all four properties into a single material to replicate the mechanical properties of tissue has been inaccessible. Herein, we present the A-g-B brush-like graft copolymer platform as a framework for fabrication of materials with independently tunable softness and firmness, capable of reaching a strength of ∼10 MPa on par with stress-supporting tissues such as blood vessel, muscle, and skin. These properties are maintained by architectural control, therefore diverse mechanical phenotypes are attainable for a variety of different chemistries. Utilizing this attribute, we demonstrate the capability of the A-g-B platform to enhance specific characteristics such as tackiness, damping, and moldability.
... Strain-stiffening represents one of nature's key defense mechanisms, which protects skin from accidental damage by malicious stretching while maintaining the initial softness 28 . By mimicking the nanofibrous structure of human skin, our hybrid ionic skin exhibits very similar behavior. ...
... The calculated fatigue threshold of hybrid ionic skin is ∼2950 J m −2 , about 40 times higher than that of the ionic matrix (∼69 J m −2 ) ( Fig. 2i and Supplementary Figs. 27,28). For comparison, we plotted the fracture energies and fatigue thresholds of a few reported anti-fatigue hydrogels and elastomers in Fig. 2j (data in Supplementary Table 2). ...
Robust ionic sensing materials that are both fatigue-resistant and self-healable like human skin are essential for soft electronics and robotics with extended service life. However, most existing self-healable artificial ionic skins produced on the basis of network reconfiguration suffer from a low fatigue threshold due to the easy fracture of low-energy amorphous polymer chains with susceptible crack propagation. Here we engineer a fatigue-free yet fully healable hybrid ionic skin toughened by a high-energy, self-healable elastic nanomesh, resembling the repairable nanofibrous interwoven structure of human skin. Such a design affords a superhigh fatigue threshold of 2950 J m⁻² while maintaining skin-like compliance, stretchability, and strain-adaptive stiffening response. Moreover, nanofiber tension-induced moisture breathing of ionic matrix leads to a record-high strain-sensing gauge factor of 66.8, far exceeding previous intrinsically stretchable ionic conductors. This concept creates opportunities for designing durable ion-conducting materials that replicate the unparalleled combinatory properties of natural skins more precisely.
... However, these materials contain high water content or expensive/toxic ionic liquids, which limits the environments in which they can be employed. Alternatively, bottlebrush elastomers (BBEs) are a class of intrinsically stretchable materials that can achieve ultra-low Young's modulus without solvents, because of their highly branched architecture consisting of polymeric side chains attached to a polymer backbone (leading to reduced entanglements in comparison to linear analogues (23)(24)(25)(26)(27)(28)). For example, polydimethylsiloxane (PDMS)-based BBEs can achieve a Young's modulus of <1 kPa, (23,29) much lower than that of commercial PDMS linear elastomers such as the Sylgard 184 with a Young's modulus of 100 kPa to 3 MPa (30, 31) (Fig. 1B). ...
Understanding biological systems and mimicking their functions require electronic tools that can interact with biological tissues with matched softness. Conductive materials that match the softness of biological tissue are thus highly demanded for the construction of ultrasoft electronics. However, the commonly employed intrinsically stretchable materials usually contain solvents that limit stability for long-term use or possess low electronic conductivity. Additionally, integrating such ultrasoft and conductive materials into electronic devices is poorly explored. This article reports a solvent-free, ultrasoft and conductive PDMS bottlebrush elastomer composite with single-wall carbon nanotubes as conductive fillers. The conductive SWCNT/BBE with a filler concentration of 0.4−0.6 wt % reveals an ultralow Young’s modulus (<11 kPa) and satisfactory conductivity (>2 S/m) as well as strong wet-adhesion property. Furthermore, we fabricate ultrasoft electronics based on laser cutting and 3D printing of conductive and non-conductive BBEs and demonstrate their potential applications in wearable sensing, soft robotics, and electrophysiological recording.
