1.

Introduction

The formation of cell-containing structures is one of the approaches of tissue engineering in creating bioequivalent tissues and organs. Generally, three-dimensional (3D) porous materials act as a base of these structures, for example, decellularized tissues^1–3 or hydrogels.^4–6 Such systems can maintain cell viability, proliferative and physiological activity, and in some cases, initiate and direct cell differentiation.^7,8 At the same time, one of the key challenges of tissue engineering is the preservation of viable cells in these constructs and transplantation of large grafts to replace extensive defects. The diffusion of oxygen and nutrients in tissues and tissue-engineering constructs is limited by distances of 100 to $200\mu \mathrm{m}$.^9,10 At larger distances, foci of necrosis often occur, and the density of living cells decreases significantly.^11,12 Vascularization of the grafts after transplantation due to the growth of the recipient’s capillaries is a slow process; therefore, graft rejection frequently occurs due to insufficient vascularization and oxygen starvation.⁶ Thus, the lack of vascularization is one of the most common causes of a rejection of the transplants of pancreatic islets.^13–15

Traditionally, diffusion limitations are overcome by perfusing the tissue-engineering constructs during the formation stage. The perfusion due to diffusion and convection helps substantially in the fight against hypoxia.¹⁶ Also, the composition of the scaffolds themselves was modified by adding components of the extracellular matrix, growth factors, and hormones that promote angiogenesis and vascularization.^7,17 The cultivation of cells under conditions of low oxygen tension before their encapsulation in scaffolds was also suggested to increase their resistance to hypoxia afterward.^18,19

Another approach to improve cell viability in 3D scaffolds, as well as to control the proliferation and differentiation of stem cells, is photobiomodulation (PBM).²⁰ This method is based on short-term exposure to low-intensity monochromatic (laser) or nonmonochromatic (LED) light in the visible and near-infrared (IR) spectral regions.^21–24 The PBM effects depend on the type of cells, their state, as well as on the wavelength, dose, and intensity of the irradiation.^23–28 Various studies had shown that light irradiation can act as an agent that directs the differentiation of cells, stimulates their survival, and activates metabolism. For example, the red 630-nm LED irradiation with an energy dose of $2.5·{10}^5\mathrm{J}/{\mathrm{m}}^2$ increased the viability of odontoblast-like cells isolated from tooth pulp.²⁹ Irradiation with the near-IR 840-nm light with an energy dose of $4\mathrm{J}/{\mathrm{cm}}^2$ stimulated the synthesis of type I collagen, and, with an energy dose of $25\mathrm{J}/{\mathrm{cm}}^2$, it improved the adhesion of neural stem cells. Moreover, the proliferation of neural stem cells immobilized in a gelatin-methacrylate matrix increased by 44% after exposure with low-intensity laser irradiation at 635-nm wavelength and an energy dose of $62.5\mathrm{J}/{\mathrm{cm}}^2$.³⁰ The authors of this study had also provided the data suggesting that such an effect can enhance cell differentiation in the late stages of cultivation. In another study,³¹ it was shown that tooth pulp stem cells immobilized in bioceramics better differentiate in the osteogenic direction after low-intensity laser irradiation (the 660-nm wavelength and a dose of 2 to $4\mathrm{J}/{\mathrm{cm}}^2$). Also, irradiation with a wavelength of 780 nm effectively stimulated osteogenic differentiation of cells and osseointegration in titanium scaffolds in vivo in a model of osteoporosis.³² Near-IR irradiation accelerates a new bone formation and osseointegration of transplanted cells in bone defects in the calvaria of rabbits.³³ And while the mechanisms of the effect of red and IR irradiation on the cell are mostly similar,³⁴ IR irradiation is considered more promising for 3D structures due to its ability to penetrate deep into tissues.^35,36 Overall, light in the red and near-IR ranges with fluences around $3\mathrm{J}/{\mathrm{cm}}^2$ was found to be the most beneficial for 3D systems.^33,37–39

The main advantages of PBM are ease of use and noninvasiveness. It is suggested that the PBM can be used as a method of cell preconditioning before transferring cells in adverse conditions.⁴⁰ However, the mechanisms and effects of irradiation on cells were studied mainly on monolayer cultures; there is almost no information on the effects of the PBM in 3D tissue-engineering structures. Optically transparent hydrogels are well suited to establish the effects of the PBM in 3D systems. Thus, a 3-mm-thick hydrogel layer of poly ($N$-vinylpyrrolidone), polyethylene glycol (PEG), agar, and water is 92% and 98% transparent for irradiation at the wavelengths of 660 and 808 nm, respectively.⁴¹

Mesenchymal stromal cells (MSCs) are on top of clinical interest because of their potential use in autologous transplantation. Currently, more than 2000 patients received autologous or culture-expanded allogeneic MSCs for the treatment of different diseases.⁴² Gingival mucosa is one of the promising sources of MSCs due to availability and the minimal invasiveness of its procurement as well as the ability of gingival mucosa wounds to heal without formation of scar.^43,44 Gingival tissue, as it is routinely discarded after resective periodontal surgery, is ideal for the isolation of the human MSCs during routine procedures under local anesthesia.⁴⁵ Human gingiva-derived MSCs produce adhesive, homogeneous, stably well-proliferating cell populations that preserve the karyotype and have pronounced anti-inflammatory and immunomodulating activity.^46,47 Gingiva-derived MSCs can differentiate effectively in the osteogenic^48,49 and myogenic directions^50,51 both in 2D monolayer cultures and in 3D culture conditions, which makes the use of gingiva-derived MSCs very promising for the formation of grafts and tissue-engineering structures for recovery of the musculoskeletal system.^7,52 MSCs from various sources are widely used in the creation of 3D tissue-engineering constructs, in particular, in combination with fibrin hydrogel, for creating a bioequivalent of the skin, cartilage, and blood vessels.^53–55 Therefore, the investigation of the effects of the PBM on such tissue-engineering constructions is of fundamental and applied importance.

The aim of this work was to test the possibility of using PBM in the red and near-IR wavelengths for tissue engineering and regenerative medicine on a model of the formation of hydrogel grafts with encapsulated MSCs from gingival mucosa. For this, we measured the physicochemical and mechanical characteristics of the hydrogels and evaluated the viability, proliferative, and metabolic activity of MSCs in 3D hydrogels of different thicknesses and concentrations of gel-forming proteins.

2.

Materials and Methods

2.4.

Photobiomodulation

2.4.1.

LED experimental set-up

The cells were irradiated with nonmonochromatic LED light of red [maximum at 633 nm, 16 nm full-width at half-maximum (FWHM)amplitude] and IR (maximum at 840 nm, 36 nm FWHM) ranges using the original apparatus LDM-07 with rectangular $5\times 11\mathrm{cm}$ LED matrices [Fig. 1(a)]. The irradiated cells were in the plate at a distance of 50 mm from the surface of the LED matrices. As a reference parameter of irradiation, we used fluence ($\mathrm{J}/{\mathrm{cm}}^2$). For both wavelengths, the fluence value was the same $2.2\mathrm{J}/{\mathrm{cm}}^2$, whereas other parameters varied due to different light sources. The characteristics of the apparatus and the exposure parameters are given in Table 2. To control the spectral composition and power of the irradiation, we used an optical fiber spectrum analyzer USB 4000 (Ocean Optics) with a wavelength range from 200 to 1100 nm and a FieldMaster power meter with a sensitive measuring head LM_10HTD (Coherent) combined with a personal computer. To reveal whether the effects of PBM were caused by light and exclude possible temperature influence, we measured a temperature using a Point Thermocouple with 0.3 mm diameter under PBM with respective parameters.

Fig. 1

(a) The matrix irradiation apparatus LDM-07 with a plate installed for irradiation with $\lambda =630\pm 8\mathrm{nm}$ light. (b) Emission spectra of red and infrared irradiators normalized to their maximum intensities. (c) Scheme of the irradiation of cells in a gel with $I_0$ intensity. (d) Transmission spectra $I/I_0$ of a 1-cm-thick fibrin gel layer with and without cells.

Table 2

Parameters of the treatment with the LDM-07 apparatus.

Light sources	Matrix of LEDs	Matrix of LEDs
Wavelength (nm)	$633\pm 8$	$840\pm 18$
Power (mW)	$160\pm 20$	$320\pm 40$
Power density ($\mathrm{mW}/{\mathrm{cm}}^2$)	$1.8\pm 0.2$	$3.6\pm 0.4$
Fluence ($\mathrm{J}/{\mathrm{cm}}^2$)	$2.2\pm 0.2$	$2.2\pm 0.2$
Energy (J)	$96\pm 10$	$96\pm 10$
Time (s)	1200	600
Number of sessions	1	1

2.4.2.

Irradiation of cell-containing gels

The cells were irradiated in two modes a day after they were encapsulated in a hydrogel. Irradiation was conducted in the dark at a temperature of 37°C for 1200 and 600 s for red and near-IR light, respectively.

The irradiation of the plates was performed in two modes at the wavelength of 633 nm for 1200 s or at the wavelength of 840 nm for 600 s, with an energy dose of $2.2\pm 0.2\mathrm{J}/{\mathrm{cm}}^2$ in both cases. A day after irradiation, the cell viability, proliferation, and mitochondrial activity were analyzed by a set of methods (PicoGreen assay, AlamarBlue assay, live/dead assay, and mitochondrial assay).

3.

Results

Despite the turbidity of the native fibrin, samples of 5:1 PEGylated fibrin were transparent. The light transmission through the modified fibrin gel was high: 96% at a wavelength of 630 nm and 99% at 840 nm [Fig. 1(d)]. Interestingly, after encapsulating cells into the gel, the resulting gel transmission did not drop but actually increased [Fig. 1(d)]. Figure 2(a) shows that the PEGylated fibrin had a flocculent structure formed by short fibers; there were uniformly distributed pores varying in diameter (0.1 to $7.2\mu \mathrm{m}$). All the gels measured were soft (the Young’s moduli $<2\mathrm{kPa}$), the average Young’s modulus values are presented in Fig. 2(b). The PEGylation led to a slight decrease of the Young’s modulus for the PEGylated fibrin gel (by 18%), whereas increase in fibrinogen concentration (from 25 to $50\mathrm{mg}/\mathrm{mL}$) led to a very pronounced increase in stiffness (by 240%) [Fig. 2(c)]. The apparent viscosity of the PEGylated gels was about twice that of the native gel.

Fig. 2

Fibrin gel characterization: (a) 5:1 PEGylated fibrin. Gel has a slightly inhomogeneous porous structure formed by short fibers; (b) mechanical properties of fibrin gels ($\text{mean}\pm \mathrm{SD}$) measured by atomic force microscopy; (c) 3D representation of Young’s modulus distributions over a $50\times 50{\mu \mathrm{m}}^2$ area mapped using the force volume mode. All gels demonstrated approximately the same level of the local heterogeneity of Young’s modulus.

The immunophenotype of the primary culture of MSCs obtained from the gingiva mucosa met the criteria for MSCs.⁶⁰ The cells used in the study expressed characteristic markers of MSCs (CD90, CD73, CD105, and CD44) and did not express hematopoietic and leukocyte markers (Table 3).

Table 3

Immunophenotype of MSCs (passage 4) from gingival mucosa.

After the encapsulation of MSCs in the gel and PBM, the effects of irradiation were analyzed by various methods for assessing cell viability and proliferation as well as by mitochondrial activity (Fig. 3). Viable cells were observed, and no significant cell death was detected in any of the samples [Figs. 4(a) and 4(b)]: the fluorescence intensity of propidium iodide was less than 1% relative to the fluorescence intensity of calcein. An assessment of the amount of DNA showed that the proliferation rate is reduced in thick gels [Fig. 4(c)], and the results of the AlamarBlue assay indicate a less active metabolism in both thick and concentrated hydrogels [Fig. 4(d)]. Thus, in hydrogels of a given thickness and concentration, the cells do not die but only switch to an inactive state, in which proliferative and metabolic activity decreases. In some cases, PBM was able to neutralize adverse conditions. The IR irradiation stimulated cell proliferation and metabolism in thick hydrogels [Figs. 4(c)–4(e)]. The results of all three tests correlated with each other: the PicoGreen assay indicated an increase in the number of cells after 840-nm irradiation by more than 2.5 times (however, the DNA content was only 60% of the control), the live/dead assay showed 20% increase, and the AlamarBlue assay showed 10% increase in metabolic activity.

Fig. 3

Experiment design. Mesenchymal stromal cells (MSCs) were isolated from gingival mucosa from the retromolar area and immunophenotyped. Bovine fibrinogen was modified with PEG-NHS for 2 h. To obtain a hydrogel with encapsulated cells, three components (MSCs, modified fibrinogen, and thrombin) were mixed in the wells of the plate and then incubated overnight in the dark at 37°C.

Fig. 4

(a) Visualization of MSCs in the hydrogel ($25\mathrm{mg}/\mathrm{mL}$) after irradiation, live/dead assay; alive cells stained with calcein (green), dead cells stained with propidium iodide (red), nuclei stained with Hoechst 33258 (blue). The scale bar is $100\mu \mathrm{m}$. (b) Magnified view, the scale bar is $100\mu \mathrm{m}$. (c) Results for PicoGreen assay. (d) Results for AlamarBlue assay. (e) Results for live/dead assay, calcein fluorescence intensity in living cells encapsulated in the hydrogel. Irradiation was conducted 1 day after encapsulation, and AlamarBlue, PicoGreen, and live/dead assays were conducted 1 day after irradiation. Hydrogel modifications in respect to fibrin concentration and gel thickness, respectively: standard: $25\mathrm{mg}/\mathrm{mL}$, 1.5 mm thick: $25\mathrm{mg}/\mathrm{mL}$, 3 mm; concentrated: $50\mathrm{mg}/\mathrm{mL}$, 1.5 mm. *$p<0.05$ relative to other datasets in the group.

The effects of PBM on cells in concentrated hydrogels manifested in a different way. The results of the AlamarBlue assay suggest that there is a tendency for a decrease of metabolic activity after irradiation [Fig. 4(d)]. Consistent with this, the data of the live/dead assay showed that the number of living cells a day after irradiation is 30% lower than that of the control [Fig. 4(e)].

Tracking of mitochondria stained with MitoTracker Green, which provides an information about mitochondrial membrane potential, is widely used to understand general mitochondrial activity.^61–65 Although the photobleaching occurs during the experiment, the relative differences still can be marked. The results of the time-lapse recording of MSCs in hydrogels showed that mitochondria of treated cells were more active and stable. The mitochondria of the encapsulated cells were also most supported by PBM with near-IR light (Fig. 5) that led to $\sim 23\%$ increase in the mitochondrial activity by the end of the experiment (5 h).

Fig. 5

Dynamics of mitochondrial activity of the MSCs encapsulated in a hydrogel ($25\mathrm{mg}/\mathrm{mL}$, 1.5 mm). High-content screening, MitoTracker Green. The control is the initial fluorescence intensity of MitoTracker Green for each of the datasets. *$p<0.05$ relative to other datasets in the group.

4.

Discussion

Fibrin hydrogel is a promising material for tissue engineering and regenerative medicine due to several advantages. A gel can be obtained from components of a patient’s blood; thus, it might be autologous.^66–68 Various 3D structures can be formed from the fibrin gel, in particular, by molding and electrospinning methods.^69,70 Moreover, the degradation rate of this hydrogel can be controlled locally, by, for example, the addition of fibrinolysis inhibitors.^71,72 Various modifications of the hydrogel can help in directing the differentiation of cells, in particular, in the angiogenic direction.^7,55

The data obtained on the gel structure agreed with the earlier obtained results.^5,73 A positive correlation between fibrin gel stiffness and fibrinogen concentration was shown in previous studies that allows the preparation of fibrin gels with desired mechanical properties in a range of low Young’s modulus values ($<10\mathrm{kPa}$).⁷⁴ The PEGylation is another factor, which affects gel stiffness by reducing it. The 5:1 PEGylated gel was also more viscous than the native one, which might be caused by smaller pore sizes. According to the poroelasticity theory,⁷⁵ the apparent viscosity is higher due to a relaxation of a solvent moving through the porous polymer network with smaller pores.

The level of light irradiation of cells in a 3D scaffold will vary due to the absorption and scattering of light in the volume of the scaffold. Due to this variation, a part of the cells may be exposed outside the “therapeutic range” of PBM. Therefore, it is important to know the distribution of light intensity in the entire matrix, which will depend on the wavelength of the light used. This information can be provided by numerical calculation involving experimental data about effective optical materials of scaffolds. The distribution of light intensities can be estimated based on the measurement of the transmission spectra of thin films.⁷⁶

Spectrophotometric measurements [Fig. 1(d)] showed that light transmission through the gel generally decreases with decreasing wavelength. In both curves corresponding to the gel without cells and the gel with the encapsulated cells, a clearly distinct region was presented in the edge of the IR range (920 to 1000 nm) associated with a local absorption peak of water. For the present study, it is of interest how quickly the red ($\lambda =633\mathrm{nm}$) and IR ($\lambda =840\mathrm{nm}$) light decays with depth upon irradiation of cells in a gel. Figure 1(d) shows that when cells in standard concentration are added to the gel, more light passes through the cuvette (the green curve lies above the blue). Moreover, in the case of a cell-containing gel of 1-cm thickness, the intensity of red light $I$ decreases to 42% of the initial value ($I_0$) and of IR light to 47% of the $I_0$.

The decrease in the light intensity occurs due to its absorption and scattering on microscopic inhomogeneities of the refractive index and density of the medium. According to Beer–Lambert’s Law, the intensity of light decreases exponentially with depth in the material:

Therefore, due to the absorption and scattering of the light in the gel, the cells will be irradiated irregularly. Exposure doses will gradually decrease for the deeper cells. In the studied gels, the irregularity of irradiation of cells in the gel with a thickness of 1.5 mm is 11% and 12%, and in the layer with a thickness of 3 mm is 20% and 23% for near-IR and red light, respectively. The difference between intensities for the red (633 nm) and near-IR (840 nm) lights is small and can be neglected.

Another important point of using PBM for 3D structures is temperature. The data obtained showed that during 600 s of irradiation with a wavelength of $\lambda =840\pm 18\mathrm{nm}$ and an irradiation intensity of $3.6\pm 0.4\mathrm{mW}/{\mathrm{cm}}^2$, the rate of heating is rather small and does not exceed $0.14\pm 0.02°\mathrm{C}$. At $\lambda =633\pm 8\mathrm{nm}$ and an irradiation intensity of $1.8\pm 0.2\mathrm{mW}/{\mathrm{cm}}^2$ for 1200 s, the degree of heating is even less than $0.09\pm 0.02°\mathrm{C}$. Thus, no local heating of the cells was observed.

Precise dosimetry and characterization of PBM parameters are crucial for comparing results. While intensity and duration of irradiation may vary from one light source to another, fluence depends only on light energy delivered to the object and reflects both the duration of the radiation and the intensity of the source in a linear manner.⁷⁷ Moreover, fluence is the most indicated parameter for articles on PBM research, and using fluence as the primal parameter makes it more reliable to analyze results.^37,78 In Refs. 31, 33, 38, 39, and 7980.–81, fluences from 2 to $4\mathrm{J}/{\mathrm{cm}}^2$ were shown as the most effective for stimulating cells within scaffolds. Based on these, we chose the fluence of $2.2\mathrm{J}/{\mathrm{cm}}^2$ for both red and near-IR light to investigate further effects.

The dependences of the fluorescence intensity (modified PicoGreen method) on the thickness of the gels [Fig. 4(c)] indicate that the activity of immobilized cells decreases with an increase in the gel thickness from 1.5 to 3.0 mm. This inhibition effect can be explained by diffusion restrictions that arise with an increase in the thickness of the scaffolds under static, nonperfused conditions. These limitations can be associated with both a lack of oxygen and a lack of nutrients. A decrease in cellular activity is a negative factor in the reconstruction of tissues and organs. To solve this problem, we proposed to stimulate cells with low-intensity irradiation with wavelengths of 633 and 840 nm. A recent study showed that blue light irradiation inhibited gingiva-derived MSCs proliferation in 2D culture, as indicated by [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide]-test (MTT), and promoted osteogenesis.⁸²

According to the results of several viability tests, changes in the physiological activity of cells in the same hydrogel samples varied greatly [Figs. 4(c)–4(e)]. Such a difference in recorded changes may be related to the sensitivity and accuracy of the methods used in these conditions. All three used assays (PicoGreen, AlamarBlue, and live/dead) are designed primarily for monolayer cell cultures. In the above experiments, a fibrin hydrogel was used as a 3D medium, which is a concentrated protein solution (5%) and acts as a turbid scattering medium. Therefore, a standard set of cytotoxicity tests might require additional calibration and optimization for 3D protein environments.

In the case of “thin gels” (thickness 1.5 mm), only a slight difference was recorded between the irradiated and unirradiated samples. In the case of gels with a thickness of 3 mm, irradiation stimulated proliferation, and this effect was especially pronounced during PBM with wavelength of 840 nm. This difference is most likely due to the specific effects of irradiation on cells. The mitochondrial respiratory chain is considered as the main target of both types of irradiation in the cell.⁸³ Absorption of light by cytochrome c oxidase leads to increasing of membrane potential, exceeded ATP production, and following fluxes of protons and calcium ions.⁸⁴ Alternative PBM mechanism involves production of a small amount of reactive oxygen species (ROS).³⁶ ROS can act as mediators in several cellular pathways including kinase pathways activating cell division.^85,86 Both of these mechanisms were shown for red and near-IR light. However, preferred paths of PBM influence on a cell may vary according to the wavelength.⁸⁷ Thus, ROS amount produced in the cells was different for red and near-IR light with equal fluencies.³⁶ Near-IR light activating cell cycle represented higher rates of ROS, which could explain more pronounced proliferation after exposure to 840 nm irradiation in the current work.

Near-IR light is more promising for tissue engineering because it is located inside the optical window and can penetrate deeper into tissue-engineered structures than red light. However, the irradiation effect was not observed in the case of thin gels with a higher concentration of fibrin ($50\mathrm{mg}/\mathrm{mL}$). Moreover, according to the results of AlamarBlue and the live/dead assays [Figs. 4(c) and 4(d)], when cells are irradiated in concentrated gels, their viability decreases. It is possible that under conditions of increased hydrogel concentration, cells may become more sensitive to stress, and thus, irradiation of the used intensities has an adverse effect.

5.

Conclusion

Hydrogels with an increase in the thickness or density decrease cell viability and their physiological activity. We have shown that it is possible to stimulate mesenchymal stem cell proliferation and metabolic activity in fibrin hydrogel using PBM. Thus, PBM can be used in tissue engineering to control cell populations immobilized in 3D scaffolds.