English

• Brain stimulation is an effective method to directly interfere with pathological neural circuits, which mainly rely on optical or electrical stimulation, to achieve the effect of treatment and research of mental diseases [1,2]. Compared with electrical stimulation, optical stimulation has the advantage of high precision without affecting surrounding non-relevant neurons, and is considered to be a new generation of important technologies for the development of neuroscience [3–6]. Optogenetic generated by combining cell biology and genetics, which modulate the activity of targeted neurons through genetic expression of photosensitive proteins, could help develop new treatments for a range of brain-related disorders [7,8].

  As is well known, the constraints of physical factors are one of the difficulties in implantable devices, The brain tissue with lower elastic modulus (in the kPa range) has high scattering and absorption properties [6,9]. Therefore, the current challenge is how to stably transmit light through complex neural tissues to target areas, which requires the implanted device to match the mechanical and chemical properties of the brain tissue, in order to avoid chronic damage and inflammatory reactions caused by implantation [10]. Some traditional silicon-based or polymer based optical fibers can be easily implanted in designated areas with excellent light guiding property, but the mismatched mechanical properties lead to tissue damage, and the wrapping of glial scars prevents them from working for a long time [11,12]. The use of flexible materials such as thermoplastic resins or elastomers is the main trend in the future [13]. In addition, another factor concerned is the balance between tissue damage caused by high modulus and implantation difficulties caused by low modulus [14].

  Hydrogel is a kind of material that possesses the nature of wet and soft, which is close to the tissues of organisms, and exhibits well biocompatibility [15–17]. Among the different hydrogel materials, hydrogel fibers both have the advantages of hydrogels and the structural superiority of fibers, thus it is easily to functionalize. By controlling the phase separation of the hydrogel network, it can realize the highly light-guiding within the hydrogel fiber, which is now named hydrogel optical fiber (HOF) [18,19]. The fabricated HOF can be used for photomedical treatment in deep tissue, such as optogenetics, photothermal therapy [18–20]. Thus, fabricating a HOF with tissue-like the mechanical modulus can meet the requirements of implanting into human tissues.

  Herein, a kind of core-sheath structured modulus self-adaptive hydrogel optical fiber (MSHOF) with adjustable mechanical properties has been fabricated by the dynamic wet-spinning technique [4,21,22]. This hydrogel fiber is artificially controlled to obtain stable signals in deep tissues for a long time, and its mechanical properties are changed through molecular design in different application scenarios. The design of the core-sheath structure [sheath: Ca-alginate; core: p(poly(ethylene glycol)diacrylate-co-N,N-dimethylacrylamide (p(PEGDA-co-DMAAM))] allows for total reflection of light at the interface, resulting in excellent light guiding property. Simultaneously the fibers fabricated possess mechanical properties with adjustable modulus (0.32–10.56 MPa) and excellent biocompatibility, which have been explored for feasibility in photogenetic stimulation and are expected to be applied for long-term neural activity modulation to treat deep neural tissue diseases. This work demonstrates significant advantages over traditional optical fibers in terms of mechanical compatibility and biocompatibility with human tissues, proving that our hydrogel optical fibers can be used in clinical experiments as efficient and controllable biomedical devices, which can continue to be designed and developed as building blocks for future photomedical treatment.

  As for the photomedicine in deep tissue, modulus is crucial to the long-term stable application of the implanted device. A large differentiation of mechanical modulus between brain/materials interface will cause foreign body reaction and inflammation, resulting in the proliferation of astrocytes and microglia, which could form glial scar tissues (thickness of about 100 µm) and cover the implant, thus seriously reducing the working stability of the implanted device [14]. Herein, the modulus-adaptive of the MSHOF is mainly divided into two aspects (Scheme 1). One is the regulation of crosslinking density, which will greatly affect the modulus of the hydrogel material. Another aspect is that when it is used as an implanted device, the MSHOF can adapt to absorb water molecules in the surrounding tissue, forming a similar state. By implanting with the MSHOF, shear micromovements of the brain caused by movement behavior of the organism can be tolerated by the low modulus of MSHOF, thus guaranteeing the safety of the biological tissue/material interface, achieving long-term modulation of neural activity.

  Scheme 1

  

  Scheme 1.  Design of MSHOF with self-adaptive modulus for modulation of neural activity.

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  The method of integrated dynamic wet spinning is used here to fabricate MSHOF as reported by our previous works [4,17], the schematic illustration of the spinning processing is shown in Fig. 1a. In brief, the solution of monomers (including PEGDA and DMAAm) is used as the core spinning layer, which was defined here as (P_XD_100-X)₅₀, P represents PEGDA, and D represents DMAAm. Between those monomers, PEGDA is used as a crosslinker, whose concentration is changed for tuning the crosslinking density of MSHOF. Na-alginate solution is used as the sheath spinning layer, and Na-alginate will rapidly ionically be crosslinked in coagulating bath (CaCl₂ solution, 2.0 wt%) once it was extruded out of the coaxial needle (Fig. 1b) [16,23]. The rapidly formed sheath hydrogel layer would bring the core spinning layer into the ultraviolet (UV) light irradiation region, and triggered the polymerization among the monomers by the initiator (I2959). The polymer network formed by the polymerization of monomers (PEGDA and DMAAm) in the core layer is shown in Fig. 1c. In Fig. S1a (Supporting information), there is no stretching vibration peak of MSHOF at the peak of around 980 cm⁻¹, the absence of C = C bonds proves that the monomer is completely cross-linked into a network. Additionally, soaking MSHOF in deuterated reagents and the absence of active hydrogen in nuclear magnetic resonance (NMR) spectroscopy also confirms this viewpoint (Fig. S1b in Supporting information) [21,22]. After confirming the stability of the fiber network, the appropriate winding speed (50–60 cm/min) is selected to collect MSHOFs, which prevents the influence of external forces on the uniformity of fibers and the coherence of spinning.

  Figure 1

  

  Figure 1.  Wet spinning process and characterization of MSHOF. (a) Schematic diagram of integrated dynamic wet spinning of MSHOF. The chemical reactions of the sheath layer (b) and core layer (c) during spinning. (d, e) The cross-sectional and side-view optical images of (P₂₀D₈₀)₉₀ in deionized water. (f) Average diameter of MSHOFs. (g) Refractive index of core spinning solutions in liquid phase. (h) A rough comparison between the MSHOF, SOF and POF. Data are presented as mean ± standard deviation (SD) (n = 5).

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  The optical fibers with different content of PEGDA were fabricated (details of spinning ratio in Table S1 in Supporting information), and named (P₁D₉₉)₅₀, (P₃D₉₇)₅₀, (P₅D₉₅)₅₀, (P₁₀D₉₀)₅₀, (P₁₅D₈₅)₅₀ and (P₂₀D₈₀)₅₀, which all have obvious core-sheath layer structure from the side-view and cross-sectional optical images (Figs. 1d, e and Fig. S1 in Supporting information). Based on the actual image, the fabricated MSHOF intuitively exhibits high transparency, indicating a high light guiding property. In addition, it can be found that the hydrogel optical fiber has a relatively uniform diameter of approximately 600–850 µm as shown in Fig. 1f. When the PEGDA content is relatively low, the network structure of the core layer is relatively loose, which leads the diameter of the hydrogel fibers would expand to a certain extent under the water. As for light guiding property, it is known that the core layer material should have a higher refractive index (RI) than the sheath layer [4,24]. Herein, the RI of the core layer hydrogel is tested as higher than 1.40 (Fig. 1g), which is larger than the sheath Ca-alginate hydrogel (RI: ~1.33) [4,17]. Such a differential RI of core and sheath materials endows the uniformity of light propagation through the hydrogel fiber [25]. Meanwhile, in a rough comparison, the fabricated MSHOFs have advantages of anti-fatigue, water content, softness, stretchability and anti-inflammatory than these properties of silica optical fiber (SOF) and polymer optical fiber (POF) (Fig. 1h).

  Scattering is an inherent property of light, which causes attenuation with depth during transmission in tissues. Therefore, optical fibers can be used to reduce light losses that are caused by tissue blocking effects. The fabricated MSHOF has an excellent light guiding property. The light propagates along the optical path of the fibers rather than scattering through the core layer, because the refractive index of the cortex is lower than that of the core layer, so total reflection of light at the interface causes relatively small losses [3,13,20].

  The light guiding property of the MSHOF is evaluated by irradiating a tip of the MSHOF with blue laser light (λ = 472 nm) (Fig. 2a), and the light attenuation, which is used for evaluating the light guiding property, is obtained by analyzing the scattered light intensity of profile photos of the light guiding (~10 cm) [4,18]. Results show that the fabricated MSHOF has low optical attenuations, which are within the range of 0.12–0.21 dB/cm (Fig. 2b and Fig. S3 in Supporting information), indicating the uniformity and excellent light guiding property of the fabricated MSHOF. It is found that the light attenuation of MSHOF presented a tendance of first increased and then decreased, this is considered that the (P₁D₉₉)₅₀ and (P₃D₉₇)₅₀ with larger diameters have smaller light attenuation values, as larger diameters lead to less contact frequency between light and the interface, resulting in less optical loss. The stable and low light attenuation indicate that the fabricated MSHOF has excellent light guiding property, and can be used for photomedical treatments such as optogenetics.

  Figure 2

  

  Figure 2.  Light propagation through the MSHOF. (a) Photos of light propagation through different MSHOF with blue light (λ = 472 nm). Scale bar: 3 cm. (b) Light attenuation calculated based on the light scattering intensity of the MSHOF from above images. Data are presented as mean ± SD (n = 5).

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  Appropriate mechanical modulus is required for implantation of hydrogel fiber into brain for long-term regulation of neural activity. The mechanical modulus that matches the brain tissue can effectively avoid implantation injury and inflammatory response, which is beneficial for the reception and transmission of optical signals [19,26,27]. Herein, we control the mechanical properties of MSHOF mainly through two different aspects. One is adjusting the crosslinking density by variating percentage of PEGDA to monomer concentration, where the PEGDA is used as crosslinkers. By this way, in the entire system, the p(PEGDA/DMAAm) in the core layer dominates the mechanical properties compared to sheath layer of Ca-alginate [22]. The specific mechanical properties (strength, modulus, elongation) of different samples are summarized in Figs. 3a and b, demonstrating that with the PEGDA content increased from 1% to 20%, the elongation of the fibers decreased from 59.34% to 22.1%, and the Young's modulus increased from 0.32 ± 0.07 MPa to 2.70 ± 0.09 MPa. This large tunable modulus range is due to the PEGDA content tuning in MSHOF, a lower PEGDA content would cause a low crosslinking density in the core layer, and a loose network structure leads to a soft property. Another way to fit the mechanical modulus of the biological soft tissue and hydrogel fiber is to control the water absorption during implantation. It can be seen from Figs. 3c and d that the mechanical modulus of the dried MSHOF is 127.8 MPa, which is higher than some soft tissue, such as liver and brain tissue. Here, we designed an experiment, in which the dried MSHOF was inserted into hydrogels with different water content (water content of 40%, 60%, 80%) to simulate the insertion in different soft tissues. Results (Figs. 3c and d) show that after inserting into the hydrogel and absorbing the water, the water content of MSHOF could reach to the value of the corresponding hydrogels that the MSHOF inserted (Fig. S4 in Supporting information). And the corresponding mechanical modulus are 10.56 ± 0.21, 3.77 ± 0.16, 1.95 ± 0.22 MPa, respectively. Based on the above two mechanisms, we can achieve modulus regulation of MSHOF in a wide range (0.32–10.56 MPa). As shown in Fig. 3e, the mechanical properties of our work are compared with various typical optical fibers, it can be intuitively seen that the performance is superior to traditional polymer based (about 10³ MPa) and silicon based optical fibers (10⁴–10⁵ MPa) [13,28,29]. Also, the modulus of MSHOF is close to many soft tissues, such as brain tissue (0.005–0.06 MPa), liver (0.05–0.25 MPa) and skin (0.8–40 MPa) [30], which shows suitability of MSHOFs for applications of photomedicine.

  Figure 3

  

  Figure 3.  Mechanical properties of the MSHOF. (a) Tensile stress-strain curves of the MSHOF with different PEGDA contents. (b) Elongation, strength and modulus of MSHOF dependent on the PEGDA contents. (c, d) Tensile stress-strain curves and its derived elongation, strength and modulus of MSHOF insert in hydrogel with different water content. (e) Comparison of mechanical properties of our work with other optical fibers and human tissues. Images of implantation of dry (f) and wet (g) hydrogel fiber into agarose gel. Scale bar: 5 mm. (h) Schematic illustration of water absorption and expansion process of dried MSHOF in brain tissue. Data are presented as mean ± SD (n = 5).

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  In addition, regarding the implantation ability of optical fibers, we used an agar model as brain tissue to simulate it. The hydrogel optical fibers cannot be implanted directly into the model (Fig. 3f), but can be easily penetrated into the model to reach the designated position in the dry state (Fig. 3g). Then fibers expand by absorbing water to restore the original bio-compatibility and mechanical properties (Fig. 3h), which have strong stability and can be used for optical signal transmission for a long time. As for swelling stability, the fibers were soaked in saline solution (0.9 wt% NaCl, simulating a humoral environment) for swelling testing. Results show that the MSHOF firstly exhibits a slight decreasing of weight and diameter, then will be constant in a value (Fig. S5 in Supporting information), indicating an excellent swelling stability during implantation. In addition, after soaking for 7 days, the shape does not change much. Considering the appropriate mechanical properties and stable swelling behavior of the MSHOF, it can be used for internal treatment to regulate neural activity by light stimulation for a long time.

  As for the applications, in vivo implant materials need to have good biocompatibility. In vitro and in vivo biocompatibility experiments show that hydrogel fibers have excellent cellular and tissular compatibility [10]. Cell counting kit-8 (CCK8) cell proliferation experiment showed that there was no statistical difference in cell proliferation activity between the hydrogel fiber group and the blank group, as shown in Figs. 4e and f, indicating that hydrogel fiber did not inhibit cell viability. The results of live/dead cell staining in Figs. 4a–d showed that the number of viable cells in the hydrogel fiber extract medium group was similar to that in the untreated group (all live cells in this experiment showed green fluorescence). It is suggested that the hydrogel fiber has good cytocompatibility in vitro.

  Figure 4

  

  Figure 4.  The biocompatibility evaluation of MSHOF with tissues. The live/dead cell staining assay (a–d) and cell proliferation activity (e, f) of NIH-3T3 cells that co-cultured with MSHOF on day 1 and 3. Data are presented as mean ± SD (n = 5). Scale bar: 100 µm. (g, h) tissue section and staining after implantation of MSHOF into skin for 7 days and 1 month. Scale bars: 2 mm and 400 µm.

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  After subcutaneous implantation of hydrogel optical fiber, major organ toxicity was detected seven days and four weeks later, blood routine function was tested seven days after implantation, respectively, and the results showed that no obvious abnormalities were found in the five viscera, and the blood routine indicators were within the normal range (Fig. S6 in Supporting information). In addition, the subcutaneous degradation of the hydrogel and the inflammation around the material were observed. The rats were dissected at 7 days and 1 month, respectively (Figs. 4g and h). The skin tissue containing the hydrogel fiber was sliced and stained, and the morphology of the hydrogel material was comparatively intact.

  Optogenetics is a new type of light-controlled cell biology technology. In recent years, optogenetics technology has developed rapidly, and its application has gradually expanded from the field of neuroscience to cell biology, pharmacology, physiology, animal behavior and other fields, and has highlighted its unique advantages in disease treatment and medical research [31–33]. At present, due to the advantages of high spatial accuracy, high spatio-temporal specificity, optogenetics has been widely used in a variety of brain diseases, hearing, vision research, and so on [34].

  However, the Young's modulus of traditional glass fiber is too different from that of brain tissue, which may lead to injury and inflammation. In order to explore the feasibility of applying our hydrogel fiber to optogenetic stimulation, a mouse primary motor cortex regulation model was constructed (Fig. 5a). All animal protocols were approved by the Animal Care and Use Committee of Donghua University and Shanghai Shengchang Biotechnology Co., Ltd. Hydrogel fiber connected with ceramic cannula and ordinary silica glass optical fiber were implanted into the cerebral cortex at the same time, which are setting as the experimental and control groups. As shown in Fig. 5b, channelrhodopsin virus (ChR) was injected into the caudal forelimb area (CFA) of primary motor cortex (CFA in M1), and hydrogel fiber was implanted and fixed at the same time. After transfection for two weeks, open field test was performed to detect the activation effect of photosensitive protein on motor cortex during light stimulation. Before light stimulation, the mice were placed in open field to explore freely to adapt to the environment. Fig. 5c shows the movement trajectory image of mice. When the light was turned off, the rats exhibited more free state of exploration. However, when the light stimulation was turned on, the mice were more excited to move with more motion state of rotation, and the average distance and speed of the movement are significantly increased (Figs. 5d and e). This may be due to the activation of motor neurons in CFA brain region by light stimulation (transporting through MSHOF), motor neurons would be activated, thus increasing the motor properties of mice, that is, their movement distance and average movement speed would increase. When the laser is turned off, the target neuron ceases its activity, and its motion state returns to that before the laser stimulation. It shows that our hydrogel fiber can effectively transmit light and has great potential in optogenetic applications.

  Figure 5

  

  Figure 5.  Application of photogenetic stimulation. (a) Construction of a motor cortex regulation model enables light and channelrhodopsin virus. (b) The target brain area for neural regulation. (c) Images of mouse movement behavior during light stimulation in field experiments. The movement distance (d) and average speed (e) of mice during optogenetic stimulation (pre-stimulus: 0–5 min, during stimulating: 5–10 min, post-stimulation: 10–15 min). Data were analyzed using ordinary t-test in Excel. ***P < 0.001.

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  In addition, we evaluated the effects of immune response and blood-brain barrier function one month following the implantation of hydrogel fiber and conventional silica glass fiber on the brain [4,35]. A marker of astrocyte activation, glial fibrillary acidic protein (GFAP), is increased in response to brain injury. GFAP expression was higher near the silica fiber implant than the hydrogel fiber implant (Figs. S7a and b in Supporting information), indicating that our hydrogel fiber implant elicited a lower immune response than the regular fiber implant. Immunoglobulin G (IgG), a common principal component of immunoglobulin in serum, is used as a marker of blood-brain barrier breach (Figs. S7a and c in Supporting information). However, IgG levels were found to be higher around silica fiber than hydrogel fiber implants. Evidence shows that hydrogel fiber has better long-term mechanical interaction with brain tissue and better biocompatibility.

  In conclusion, a kind of core-sheath structure hydrogel optical fiber with adjustable mechanical properties was designed and continuously prepared by the dynamic wet-spinning technique. The research focus of this experiment is the molecular design of the core layer, and the performance of the hydrogel fiber is regulated by the proportion of PEGDA in the monomer. The fabricated MSHOF has adjustable mechanical properties (0.32–10.56 MPa) and excellent light guiding property (0.12–0.21 dB/cm) along with good biocompatibility (low biological toxicity and inflammatory response after implantation). In addition, this study demonstrated the potential application of MSHOF in long-term modulation of neural activity by implanting it into the motor cortex of mice for photogenetic simulation. Therefore, the molecular design and structural regulation in this paper provide new perspectives and application directions for the continuous preparation of hydrogel fibers with controllable mechanical properties, which has great research prospects in neuroscience, intelligent optical medicine and other fields.

  Declaration of competing interest

  The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

  CRediT authorship contribution statement

  Guoyin Chen: Writing – review & editing, Writing – original draft, Validation, Methodology, Investigation, Data curation. Siming Xu: Investigation, Data curation. Zeqi Zhang: Writing – original draft, Data curation. Ying Guo: Investigation. Jiahao Zheng: Data curation. Jialei Yang: Investigation. Jie Pan: Writing – original draft, Conceptualization. Kai Hou: Writing – original draft, Conceptualization. Meifang Zhu: Conceptualization.

  Acknowledgments

  This work is supported by the National Key Research and Development Program of China (Nos. 2021YFA1201302 and 2021YFA1201300); the National Natural Science Foundation of China (Nos. 52303033, 52173029); Shanghai Sailing Program (No. 23YF1400400); the Natural Science Foundation of Shanghai (No. 21ZR1400500).

  Supplementary materials

  Supplementary material associated with this article can be found, in the online version, at doi:10.1016/j.cclet.2024.110440.

  
  
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• 

  Scheme 1  Design of MSHOF with self-adaptive modulus for modulation of neural activity.

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  Figure 1  Wet spinning process and characterization of MSHOF. (a) Schematic diagram of integrated dynamic wet spinning of MSHOF. The chemical reactions of the sheath layer (b) and core layer (c) during spinning. (d, e) The cross-sectional and side-view optical images of (P₂₀D₈₀)₉₀ in deionized water. (f) Average diameter of MSHOFs. (g) Refractive index of core spinning solutions in liquid phase. (h) A rough comparison between the MSHOF, SOF and POF. Data are presented as mean ± standard deviation (SD) (n = 5).

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  Figure 2  Light propagation through the MSHOF. (a) Photos of light propagation through different MSHOF with blue light (λ = 472 nm). Scale bar: 3 cm. (b) Light attenuation calculated based on the light scattering intensity of the MSHOF from above images. Data are presented as mean ± SD (n = 5).

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  Figure 3  Mechanical properties of the MSHOF. (a) Tensile stress-strain curves of the MSHOF with different PEGDA contents. (b) Elongation, strength and modulus of MSHOF dependent on the PEGDA contents. (c, d) Tensile stress-strain curves and its derived elongation, strength and modulus of MSHOF insert in hydrogel with different water content. (e) Comparison of mechanical properties of our work with other optical fibers and human tissues. Images of implantation of dry (f) and wet (g) hydrogel fiber into agarose gel. Scale bar: 5 mm. (h) Schematic illustration of water absorption and expansion process of dried MSHOF in brain tissue. Data are presented as mean ± SD (n = 5).

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  Figure 4  The biocompatibility evaluation of MSHOF with tissues. The live/dead cell staining assay (a–d) and cell proliferation activity (e, f) of NIH-3T3 cells that co-cultured with MSHOF on day 1 and 3. Data are presented as mean ± SD (n = 5). Scale bar: 100 µm. (g, h) tissue section and staining after implantation of MSHOF into skin for 7 days and 1 month. Scale bars: 2 mm and 400 µm.

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  Figure 5  Application of photogenetic stimulation. (a) Construction of a motor cortex regulation model enables light and channelrhodopsin virus. (b) The target brain area for neural regulation. (c) Images of mouse movement behavior during light stimulation in field experiments. The movement distance (d) and average speed (e) of mice during optogenetic stimulation (pre-stimulus: 0–5 min, during stimulating: 5–10 min, post-stimulation: 10–15 min). Data were analyzed using ordinary t-test in Excel. ***P < 0.001.

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