Hydrogels are very attractive materials for biomedical applications—they can be composed of biologically relevant and biocompatible polymers. Their physico-chemical properties can be controlled and hydrogels have large amounts of water that can work as space to host molecules of interest. Hydrogels can be used for encapsulation of biologically-relevant materials, to form injectable forms [1
], as bone defect fillers, for surface functionalization via coating and formation of 3D gel matrices which can support tissue and organ engineering, i.e., bioprinting [3
The injectable materials approach means that outside of the body, the material is liquid but solidifies in situ. Thereby, the defect is filled since the injected material can adopt the shape and size of a defect. This way, introduction is minimally invasive causing lesser discomfort for a patient [4
]. Such injectable hydrogels can be built based on proteins and polysaccharides [5
]. In addition, bone filling materials often include bone cement consisting of calcium phosphates [6
]. Hydrogel surface functionalization or bioprinting provides a special biocoating that could serve as a 2D or 3D template for cell adhesion [7
], align cells [8
], or encapsulate and release functional biomolecules like alkaline phosphatase [9
]. Various polymers can be used to construct hydrogels for tissue engineering [10
]. However, a more interesting approach is using polymers with ionic cross-linkage mechanisms, like alginate and gellan gum, due to easy control of gelation and a chemically “mild” method of gel preparation.
Various approaches have been developed to overcome the mechanical weakness of hydrogels, for example, crosslinking. However, this includes the use of chemical agents. Hybrid materials and interfaces [11
] represent an attractive alternative to chemical crosslinking and have been applied in various nanomedicine [12
] applications. For example, nanoparticles [14
] have been identified as key enabling components for the construction of robust polymeric coatings with advanced functionalities including, for example, laser activation [15
]. Carbone nanotubes (CNTs) represent a special class of carbon-based materials; they have been used for different applications including supercapacitors, photovoltaics, photodiodes, sensors [16
]. Indeed, adding carbon-based structures offers advantages because unique properties of carbon materials such as high conductivity, robustness, possibilities of biochip functionalization can be used [17
]. Earlier, it was reported that CNTs enhance the mechanical properties of freely suspended polymeric microcapsules [18
A wide range of applications in biomedicine has also been identified for materials based on CNTs and hydrogels including tissue engineering, drug delivery, theranostics, sensing, and biosensing [19
]. However, dispersing carbon-based structures represents a particular challenge. With regard to this, oxidation plays an important role [22
]. The following acid solutions have been often used for oxidation of CNTs: HNO3
]. In the oxidation, thermal processes were identified as an essential constituent [25
]. However, the possibility of using particularly loose and stand-alone nanotube for bioapplication may be hazardous, which requires additional research [26
The low tensile strength of many hydrogels limits their use for effective cell growth. Several strategies are available to increase the mechanical strength of the hydrogel. For example, it is possible to design hybrid materials [11
] by incorporating various particles or networks [27
], for example, microcapsules [28
], bioglass particles [29
], hydroxyapatite particles [30
] calcium, and magnesium carbonate particles [32
]. In this case, it is identified that the cell senses the environment and literally “grabs” hard particles at surfaces [33
]. An overview of various interfaces allows us to conclude that many particles incorporated into polymeric coatings can promote cell adhesion [36
]. Carbon-based structures represent interesting alternatives for the functionalization of polymeric surfaces, but their full potential remains to be explored in more detail.
In this work, we demonstrate that adding CNTs to rather soft gellan gum (GG) hydrogels transforms these cell “unfriendly” (repelling) coatings into a very effective cell culture platform. Analysis of the mesh sizes and structure of such hybrid coatings is performed using scanning electron microscopy (SEM) and atomic force microscopy (AFM). Also, mechanical properties of CNT-GG hydrogels are investigated at both macro- and nano-scales revealing that the addition of CNTs results in the formation of very robust and mechanically stiff coatings. Implications for cell culture growth is also discussed here. The most essential parameter enabling effective cell growth on soft hydrogel structures—the minimum CNT concentration—have been identified.
2. Materials and Methods
2.1. Hydrogel Synthesis
GG hydrogels with loaded CNTs (GG-CNTs) were formed by mixing various concentrations 0.5, 1 and 1.5% w/v (5, 10, and 15 mg/mL) of CNTs (Sigma-Aldrich, can. No. 677248-5G) with 1% w/v (10 mg/mL) Gelzan (GG CM, product no. G1910, “Low-Acyl”, 200−300 kDa) [39
]. Solutions were mixed by vortex for 5 min and US bath (Digital 10 P, Bandelin SONOREX, Germany) at HF-Frequency 35 kHz for 15 min at room temperature. After that GG-CNTs solution was crosslinked by CaCl2
(0.3 M) for 20 min. After this step, hydrogels are washed twice with water to remove unreacted ions.
2.2. Scanning Electron Microscopy
GG-CNTs hydrogels were characterized by scanning electron microscopy (SEM). Hydrogel cylinders with 5 mm diameter were cut by biopsy and transferred to an aluminum stage covered with double-sided carbon tape. Before SEM, the hydrogels were dried by a lyophilizer and coated with a 15 nm thick gold layer (Bal-tec SCD050 Sputter Coater, MA, USA). The measurement was performed with JSM-T330A from JEOL at the operating voltage of 25 kV with secondary electrons.
2.3. Mechanical Tests Using a Universal Testing Machine
Mechanical stiffness of gels has been performed using a Universal Test Machine, LS1 (1 kN) Material Tester from Lloyd Instruments, Inc. (Ametek, PA, USA). A 50 N load cell was used for making a 2 mm indentation in samples with the diameter of the tip of 10 mm (the preload was set to 0.05 N).
2.4. Atomic Force Microscopy
The AFM data were acquired using a Nanowizard 4 Atomic Force Microscope (Bruker/JPK Instruments, Berlin, Germany) operated in the quantitative imaging (QI) mode (in liquid). All maps were obtained with QP-BioAC−CI probes (Nanosensors, Neuchatel, Switzerland), using the cantilever which had a nominal frequency of 50 kHz and a force constant of 0.1 N/m (calibrated in contact-free mode). Force maps were collected using a set point of 3 nN at 1.6 ms per pixel, with a Z length of 0.2 μm and a tip speed of 125 μm/s. The force and height maps were 25 μm × 25 μm. To measure mechanical and topographical properties, AFM was utilized with the DNP-s10 cantilever (Bruker, MA, USA).
To gain information on the viscoelastic properties of hydrogels, microrheological measurements were performed using DNP-S10 chips (nominal frequency of 50 kHz and a force constant of 0.1 N/m). In these measurements, the force reaction toward small amplitude oscillating forces (at low frequencies) that are applied at the surface are analyzed. The oscillations in the experiments were performed with an amplitude of 50 nm at frequencies ranging between 10 and 200 Hz (in the following steps: 10, 20, 50, 75, 100, 150, 200 Hz). Calculations of the storage and the loss moduli were performed in JPK data processing software, based on calculations adapted from a previous paper on microrheological measurements on live cells [40
]. To accommodate for influences originating from the cantilever’s geometry, first, the deviation from 90° phase shift in the liquid environment and the hydrodynamic drag coefficient was calculated and incorporated into the measurements. Fitting the microrheological data to the soft glassy rheology model was performed by non-linear fitting in OriginPRO 2020 [41
2.6. Osteoblasts Cultivation
Osteoblastic MC3T3-E1 cells were cultured in MEM-alpha glutaMAX-1 (cat. no. 32561-029) supplemented with 10% FBS, and 100 μg/mL penicillin/streptomycin. The media were replaced every three days, and the cells were maintained in a humidified incubator at 5% CO2 and 37 °C (Innova CO-170, New Brunswick Scientific, NJ, USA).
2.7. Fluorescence Microscopy
To estimate cell adhesion and proliferation on the surface of the prepared samples, viable cells were visualized by a fluorescence microscope using a microscope Nikon TI (Nikon, Japan) with Objective 10× and appropriate filters. MC3T3-E1 cells were seeded on the sample surfaces with an area of 0.31 cm2 at a cell density of 10 × 103/sample and incubated for 1 and 3 days. Afterward, cells were stained with Calcein AM. Samples were washed two times in PBS to remove the non-adhered cells. The number of cells was calculated from snapshots of three random zones for three replicates of the samples.
2.8. Cell Viability Test
The effects of samples on MC3T3-E1 cells were determined by AlamarBlue (ThermoFisher Scientific; Cat. No DAL1025). Samples were lowered to the bottom of a 96-well plate. MC3T3-E1 cells were seeded into 96-well on hydrogel samples at a cell density of 10 × 103/well in the culture medium and incubated 1 and 3 days at 37 °C under 5% CO2. In the last step, 10 μL of fluorescence dye (AlamarBlue) was added to each well, and the fluorescent (540/610 nm) intensity was measured by a spectrophotometer (Infinite F200 PRO, Tecan, Switzerland).