Injectable Non-leaching Tissue-mimetic Bottlebrush Elastomers: A New Platform for Advancing Reconstructive Surgery

Current materials used in biomedical devices do not match tissue’s mechanical properties and leach various chemicals into the body. These deficiencies pose significant health risks that are further exacerbated by invasive implantation procedures. Herein, we leverage the brush-like polymer architecture to design and administer minimally invasive injectable elastomers that cure in vivo into leachable-free implants with mechanical properties matching the surrounding tissue. This strategy allows tuning curing time from minutes to hours, which empowers a broad range of biomedical applications from rapid wound sealing to time-intensive reconstructive surgery. These injectable elastomers support in vitro cell proliferation, while also demonstrating in vivo implant integrity with a mild inflammatory response and minimal fibrotic encapsulation.

solvent-free injectable implants that do not leach into the body and precisely replicate the deformation response of soft biological tissues remains elusive.
Gel's inability to replicate tissue mechanics is due to the inherent flexibility of polymer network strands, which impart softness at the expense of reducing firmness. 19,20 As such, state-of-the-art implants such as for body reconstruction are soft, but not firm. This hurdle was resolved by utilizing the brush-like polymer architecture where densely grafted side-chains concurrently dilute and stiffen network strands to enable elastomers with enhanced softness and firmness. [19][20][21][22][23] Independently controlling these mechanical characteristics without adding solvent as a mechanical regulator 20 allows mimicking the stress-strain response of various tissues ranging from supersoft brain tissue to tough skin. 24,25 However, in vivo application of this technology is severely limited as such elastomer synthesis typically involves solvent and crosslinking schemes that requires hazardous stimuli such as temperature and UV light. 20,24,26 To mitigate these issues, we exploit two vital traits of the brush architecture: (i) compact molecular conformation providing low melt viscosity and (ii) a myriad of chain ends apt for functionalization. This enables solvent-free in vivo injection of reactive bottlebrush melts to yield non-leachable elastomers that match the mechanics of surrounding tissue. Depending on the targeted application, the developed methodology allows fine-tuning both the Young's modulus from 10 2 of adipose tissue to 10 5 Pa of skin and the gelation time from hours to minutes to match the duration of various surgical procedures. Given the design-by-architecture approach, this technology can be translated into a broad range of chemical compositions to satisfy specific functionality and biocompatibility requirements of various applications including reconstructive surgery, regenerative medicine, drug-delivery, soft robotics, and wearable diagnostics.
Specifically, injectable elastomers are an attractive alternative to invasive deployment of reconstructive implants as they offer improved patient comfort, reduced costs, faster recovery, and minimal surgical and post-surgical complications.

Concept and synthesis of injectable solvent-free elastomers
A dual syringe formulation consists of two reactive components: (i) a melt of bottlebrushes with functionalized side chains, and (ii) a difunctional crosslinker (Fig. 1a). Upon co-injecting, the functionalized bottlebrush and crosslinker mixture spontaneously reacts to yield a super-soft elastomer ( Fig. 1b) with tunable mechanical properties by augmenting their stoichiometry as discussed below. Given the large size of bottlebrush macromolecules, a minuscule fraction of crosslinking moieties (0.02 mol%) is required to achieve a fully conjugated network, which minimizes uncontrolled reaction with surrounding tissue and precludes polymerization-induced shrinkage. To adjust the gelation time ( ), a broad range of crosslinking chemistries has been considered including isocyanate/hydroxyl, isocyanate/amine, aldehyde/amine, alkyne/azide, and diene/dienophile (Fig. 1c). [27][28][29][30][31] In this study, we primarily explore two systems for fast ( ~ minutes) and slow ( ~ hours) gelation rates respectively using isocyanate:hydroxyl (NCO:OH) and isocyanate:amine (NCO:NH2) coupling (Fig. 1d), yet other crosslinking schemes such as Diels-Alder chemistry can be considered if longer curing times are desired. 32 Fine-tuning is also achieved by manipulating the fraction of functionalized chain ends, temperature, and catalyst concentration as discussed below. In all cases, solvent-free injection is empowered by a significant reduction of brush melt viscosity relative to linear polymers with identical molecular weight due to limited overlap and entanglement of bottlebrush macromolecules (Fig. 1e). 33 Additional decreases in viscosity can be achieved by using more complex architectures such as star-like bottlebrush melts (Fig. 1e).
To enable injectable materials with tunable mechanics, adjustable curing rate, and an enhanced biocompatibility profile, it is imperative to design brushes with reactive moieties that meet critical criteria including: (i) targeted yet small fractions of functionalized reactive side chain ends with (ii) broad post-functionalization potential that are (iii) randomly dispersed throughout the brush, and which (iv) do not interact until curing. Although these features are readily programmed individually, taken together, they represent a significant synthetic challenge in respect to affording the desired mechanical properties, curing time and biocompatibility profiles.  1), which highlights the potential for future tailored functional brushes for honing curing times and biocompatibility.  15).

Controlling gelation time of injectable elastomers
Gelation is monitored by rheology, which identifies the crossover time between the storage (G′) and loss (G″) moduli at 37˚C (Fig. 1d). Within a given crosslinking scheme (e.g., NCO:OH), a combination of stoichiometry and temperature allows tuning gelation time ( ) within more than two orders of magnitude as demonstrated by increasing by simultaneously decreasing crosslinker concentration ( Fig. 2a,b) and temperature (Fig. 2c,d), i.e. 1:1 at 50C versus 1:8 at 20C.
Similarly, switching from OH to NH2 functionalization decreases from hours to the minutes ( Fig. 1d), which can be fine-tuned in the future by mixing OH-and NH2-terminated side chains into brushes. Overall, the injectable technology contains a toolbox of architectural and chemical parameters to enable broad tuning of cure time to cover a significant portion of biomedical applications. However, it is important to note that tuning crosslinker concentration inadvertently augments both the curing time and mechanical properties. To decouple and 0 at a constant = 37℃, we prepared NCO:OH injectable formulations with different catalyst concentrations, which allows varying curing duration at a constant modulus of a fully cured elastomer (Fig. 2e).
These decoupling efforts can be also explored through additional crosslinking chemistries (   Fig. 18 and 19).

Tissue-mimetic mechanical properties
Mechanical properties of fully cured elastomers were evaluated by uniaxial tensile tests using the following equation of state relating true stress with network elongation ratio = / 0 from its initial 0 to deformed size : This stress-elongation relationship has been validated for various synthetic and biological polymer networks 20,25,34 and is described by two mechanical characteristics: Young's modulus 0 (eq 2) and firmness parameter (eq 3): where is the structural modulus controlled by crosslink density. The firmness parameter 0 < < 1 characterizes network's strain-stiffening behavior described by extensibility of network strands from their initial end-to-end distance to the fully extended contour length .
Therefore, the Young's modulus 0 of a polymer network depends not only on its crosslink density, but also on initial conformation of network strands (~〈 2 〉).    (Fig. 3d), where -Boltzmann constant,absolute temperature,monomer length, andmonomer volume, 23 which reinforces the well-defined structure of injected polymer networks. Lastly, the mechanical properties are "invariant" with respect to crosslinking chemistry highlighting the flexibility of the injectable platform chemistry to only controlling curing duration and final product biocompatibility.  (Fig. 3), injectable dynamic elastomers based on Diels-Alder chemistry ( Supplementary Fig. 19), and injectable photocurable elastomers ( Supplementary Fig. 20). 1) Injected ratio. Degrees of polymerization (DP) of 2) side-chains and 3) backbone of random polydimethylsiloxane-poly(ethylene glycol) (PDMS-r-PEG) bottlebrush macromolecules prior to crosslinking determined by 1 H-NMR. 4) Nominal DP of the backbone strand between cross-links.
As mentioned above, injectable elastomer [400,14] demonstrate nearly identical softness and firmness with a silicone gel (~30 wt% gel fraction) extracted from a commercial breast implant ( Fig. 3b), yet the solvent-free elastomers are more resilient and demonstrate significantly higher elastic deformation prior to fracture ( ). To further demonstrate the adequate mechanics of injectable elastomers, we conducted a texture profile analysis (TPA), whereby cylindrical samples are subjected to cyclic compressions at different deformations (Fig. 3e). From the TPA profiles, we evaluate several industrially relevant mechanical characteristics such as springiness, resilience, and cohesiveness ( Supplementary Fig. 21) that favorably compare the solvent-free injectable elastomers with a commercial gel containing ~70% of liquid fraction (Fig. 3f).

Non-leachability and cytocompatibility
Gel-based implants perpetually leach various chemicals such as diluents, catalysts, and ligands into the body over time and upon deformation, which represents a significant long-term health concern. 10,15,35 This is readily observed by silicone gels leaching onto a paper towel (Fig. 4a), which is quantitatively corroborated by aqueous extraction of the sol fraction in time-resolved 1 H-NMR (Fig. 4b) contrary to our non-leaching injectable elastomers. To further demonstrate the significance of leachable-free compositions , we compare cytotoxicity 36 and cell proliferation between gels and injectable elastomers. 37,38 Cytotoxicity tests are performed according to ISO 10993-5 for the aqueous extractions (Fig. 4c) with a NIH/3T3 fibroblast viability above 90% when exposed to extracts from the injectable formulations after 24 hours (Fig. 4c), while extracts from commercial silicone gel implants show significantly diminished viability of 40-60%.
Further, the proliferation of NIH/3T3 fibroblasts is analyzed by measuring the total DNA content of cultured fibroblasts. The total extracted DNA from cultured cells on elastomer surfaces confirm increasing cell count over two weeks for each injectable formulation (Fig. 4d). This is visually confirmed by time-resolved fluorescence imaging (Fig. 4e), which affirms the injectable elastomer formulations as viable biocompatible materials.

In vivo implantation of injectable elastomers
Traditionally, encapsulation of silicone gels within a stiff impermeable shell has been employed to control leaching rate. 10 However, many reports show limited improvement as the shell material is permeable to small molecules 8 and is significantly stiffer than surrounding tissue instigating capsular growth. 6 Therefore, we conduct in vivo assessment of our injectable elastomers using animal models subjected to both subcutaneous and intramuscular implantation ( Fig. 5a). In each case, explanted samples are well tolerated, with no clinical evidence of inflammatory response in surrounding tissues. In the subcutaneous explants, a thin translucent layer of encapsulating connective tissue is observed, which is significantly thicker around silicone gels. In muscle tissue, the injectable samples are fully intact and can be thoroughly explanted in contrast to the disfigured and partially fragmented silicone gels (Fig. 5a). According to the Hematoxylin-Eosin and Mallory Trichrome staining overview, the injectable elastomer capsule does not contain multinucleated foreign body giant cells at any stage and does not contain lymphocytes, leucocytes, macrophages on later stages, suggesting the implanted materials preclude chronic inflammation and are sufficiently inert (Fig. 5b, c). The capsular thickness of the fibrous layer was quantified by morphometric image analysis on the Mallory's trichrome stained slides (Fig. 5c). The injectable elastomer samples display significantly lower capsular thickness compared to silicone gels at 1, 4 and 12 weeks (Fig. 5d), which may be ascribed to both the lack of leaching into the animal body and their tissue-matching softness.

Conclusion and outlook
Soft tissues demonstrate a very distinct response to deformation: they are soft when touched, yet rapidly stiffen upon deformation. Synthetic replication and in vivo implementation of this duality is imperative for body fillers and reconstruction of various tissues after disease and injuries including HIV-associated lipoatrophy, mastectomy or lumpectomy, burns, tumor removal, and general bodily injuries. These traumas not only represent physical concerns such as severe sitting pain after loss of adipose tissue in the buttocks, 39 but also psychological stresses after mastectomy 40,41 and discrimination after loss of tissue in the face. 42,43 Although current implant strategies attempt to address these issues, they also introduce a host of additional undesired consequences. For instance, current devices only replicate tissue softness by employing various kinds of gels, i.e. by diluting polymer networks with liquids, which continually leach into the body as commonly experienced by women who undergo breast reconstruction with silicone (0.06 mg/liter) detected in their breast milk. 44 Furthermore, implant's significant firmness mismatch with surrounding tissue creates additional health risks and psychological issues due to capsular contracture and disfigurement, which requires future invasive explanation surgeries. To mitigate surgical complications, different injectable technologies have been introduced such as injecting polyacrylamide microgels (PAAG). Although initially thought to be safe, PAAGs are now banned in most countries due to substantial evidence associating PAAG with infection, glandular atrophy, fibrosis, inflammation, and palpable scleroma formation after water fraction absorption of migrated microgel fragments throughout the body. 8,45,46 Given these outlined circumstances, we believe that our minimally invasive injectable, non-leaching, tissue-mimetic, and biocompatible elastomer platform will advance various biomedical device applications.
Architecturally tailored brush based mesoblocks augmented with functionalized side chains enables both tunable curing time and tissue-like mechanical properties of fully cured materials.
The formulations contain no solvent and does not leach, while bottlebrush architecture of chains enables injectability due to significantly lower brush melt viscosity compared to linear chains of similar molecular weight. These features coordinate to bestow sufficient biocompatibility as demonstrated both in vitro and in vivo. Although this technology promises to revolutionize reconstructive implantation procedures, future work will be aimed to expand the crosslinking schemes and architectural landscape to independently control softness, firmness, and curing duration. The design-by-architecture approach is adaptable to any chemistry, which will allow future expansion of this platform to other commodity polymers such as polyolefins, polyacrylates, and polyesters. Furthermore, side chain functionalization opens many opportunities for precision engineering of alternative applications such as tissue adhesives, and coating of implanted medical devices to enhance biocompatibility and performance. 38 Last but not least, the devised injectable platform is readily applicable to fabricate soft medical implants with tissue-mimetic mechanics via additive manufacturing techniques. Synthesis of azide-terminated macromonomer. The following procedure was performed to synthesize amine-terminated PEGMA macromonomer. A 100 ml round-bottom flask equipped with a magnetic stir bar was charged with 10 g PEGMA, 50 ml DCM, and 2.5 g TEA, sealed and then placed in an ice bath. Subsequently, 2.5 g methanesulfonyl chloride was added drop-wise to the mixture using a syringe pump, and reaction was stirred overnight. The resultant solution was passed through column for purification, and then dried. The obtained PEG derivate along with 50 ml DMF and 3 g sodium azide were charged into a 100 ml round-bottom flask equipped with a magnetic stir bar. The reaction was stirred for 24 h at room temperature. The mixture was centrifuged to remove excess salt, dried and then azide-terminated PEGMA macromonomer was extracted by dissolving in DCM followed by washing with water. 1 H-NMR spectrums of PEG macromonomer functionalization at different stages are shown in Supplementary Fig. 6.

Synthesis of amine-functionalized brushes.
A similar method as described above was followed to synthesize brush polymers using the PDMSMA and azide-terminated PEGMA macromonomers. After purification of the brushes, they were dissolved in anhydrous THF, reacted with excess THPP for 24 h, and then water was added to the mixture. Finally, the aminefunctionalized brushes were purified via passing through column, and then dried for further use. 1 H-NMR spectra of PDMS-r-PEG.N3 and PDMS-r-PEG.NH2 brush copolymers are displayed in Supplementary Fig. 7.

Synthesis of polydimethylsiloxane diisocyanate crosslinker.
In order to synthesize PDMS macromolecular crosslinker, a 100 ml round-bottom flask equipped with a magnetic stir bar was charged with 10 g IPDI, 50 ml anhydrous DCM and sealed. Afterward, 5 g DMS-A15 dissolved in 10 ml anhydrous DCM was added drop-wise to the mixture using a syringe pump.
Subsequently, the resulting PDMS diisocyanate crosslinker was purified to remove excess IPDI and dried for further use. Supplementary Fig. 8  injectable elastomer, PDMS-r-PEG brushes were mixed with predetermined amount of PDMS diisocyanate crosslinker to reach predetermined crosslink density. Supplementary Fig. 13 demonstrates administration and handling of an injectable elastomer by means of a double syringe system.

Synthesis of injectable dynamic tissue-mimetic elastomers.
Reversible Diels-Alder chemistry was used to prepare injectable dynamic tissue-mimetic elastomers from mixtures of functionalized brushes and a difunctional crosslinker (Supplementary Fig. 18 and 19). To  Fig. 19).  Table 1.

Synthesis of injectable
Texture profile analysis. In order to examine how injectable elastomers behave when deformed, texture profile analysis (TPA) was performed using an RSA-G2 DMA (TA Instruments) in compression mode. Disk-shaped samples with 8 mm diameter were compressed twice, and their behavior at different strain ratios of 20, 50, and 70% was monitored. TPA parameters (springiness, resilience, and cohesiveness) were measured based on force-time curves.
Elastomers bleed (leachability) tests. In order to monitor the leachability of injectable elastomers in comparison with silicone gels used in commercial breast implants, elastomers were immersed in aqueous medium and extracted samples were monitored using 1 H-NMR at different time intervals over one month (Supplementary Fig. 9). In order to quantify the leachable fraction from three types of commercial silicone gels and our injectable elastomer after one month, the extracted samples were freeze-dried, and their mass was measured, and reported based on sample weight ( Supplementary Fig. 10). Furthermore, to visualize the leachable diluent fraction from a commercial silicone gel in comparison with the injectable elastomers, bulk samples were placed on a paper substrate and monitored over time ( Supplementary Fig. 11).  and embedded in paraffin. The 5 µm paraffine slides were sectioned using a microtome. The slides were then stained with hematoxylin and eosin (H&E) after deparaffinization and rehydrated in a graded ethanol solution series (100%, 90%, and 70% ethanol, 5 min each; dH2O for 10 min). For visualization of connective tissue, the Mallory's trichrome (Bio-Optica) was used. Further, microscopic analysis was carried out using a Leica DM750 light microscope (Leica), and digital images were recorded using an ICC 50 camera (Leica).