Functional magnetic ferrogels: From biosensors to regenerative medicine

A hydrogel is a solution (typically in water) of flexible polymer chains bonded to each other in a finite number of points. They form a 3D cross-linked polymeric network with chain sections between crosslinks containing above 50 monomer units. If the polymer chains carry an electric charge, the gel is denominated “polyelectrolyte” – an equivalent number of oppositely charged counter ions corresponds to the gel medium.^1,2 The physical properties of a gel are similar to the solution of a polymer in a liquid. There is a difference, however: a gel retains its shape and has elasticity. A significant increase in the deformation level under the application of a magnetic field can be achieved for magnetic ferrogel (FG), i.e. a multicomponent system, consisting of magnetic nanoparticles (MNPs) spread over polymeric network swollen in liquid.³ FG can be prepared by incorporating magnetic nanoparticles or even magnetic wires into chemically cross-linked hydrogels.^4–6 Nowadays a large variety of FG can be synthesized. Fig. 1 shows some typical examples.

There are different reasons for the interest in biomedical applications of gels and ferrogels. Polyelectrolyte gels and FG are attracting special interest in the area as biomimetic materials. Many features of biological tissues vary from one biological system to another especially in the case of tumor tissues.⁷ At the first stage of many biomedical studies, the substitution of biological samples by synthetic hydrogels or FG to mimic the main properties of living tissues makes sense. Synthetic hydrogels can be used as simplified models of the cytoskeleton and in addition do not require a strict protocol for testing and safekeeping.

Fig. 2 summarizes some selected prospects for the practical use of ferrogels: controlled drug release, biocompatible valves and actuators controlled by a magnetic field, magnetic biosensors, matrices for growing cells and tissues including conditions of growing in a magnetic field. Recently, we reported on the possibility of cultivating human dermal fibroblast cell cultures,⁸ which also might be useful for regenerative medicine. As in previous tests with magnetic biosensing systems, polyacrylamide-based FG with embedded magnetic nanoparticles of iron oxide were used for cultivating human dermal fibroblast cell cultures. Biocompatible γ-Fe₂O₃ MNPs were obtained by the laser target evaporation technique (LTE),⁹ which provides large batches greatly requested in nanomedicine.⁷ The above-mentioned approach of cell cultivation onto gel/FG surfaces might be useful for regenerative medicine, drug delivery and biosensor applications opening a new combination of magnetic field assisted delivery, positioning control and post operation monitoring. In-tissue embedded MNP detection is one of the most sought-after biomedical applications. In previous publications we reported on a magnetic biosensor prototype with a magnetic multilayered sensitive element for the detection of mesoscopic distributions of nanoparticles using the magnetoimpedance (MI) effect.¹⁰ MI responses were measured with and without a polyacrylamide FG layer in order to evaluate stray fields of embedded MNPs. For the development of a new generation of drug delivery systems for multifunctional devices a further increase of the sensitivity of the magnetic field sensor is necessary.

In this work, we describe our experience developing functional magnetic polyacrylamide ferrogels with examples in the field of regenerative medicine and magnetic biosensing using iron oxide γ-Fe₂O₃ magnetic nanoparticles produced by LTE.

Iron oxide MNPs were prepared by the LTE method (more details of their characterization can be found elsewhere¹¹). Polyacrylamide (PAAm) FG with a network density of 1:100 were synthesized by radical polymerization of acrylamide in a stable aqueous suspension of iron oxide MNPs. Concentrations of 2.4%, 2.1%, 1.3% and 1.0% of MNPs by weight were obtained. More details about the FG preparations can be found elsewhere.⁹ Magnetoactive epoxy composites with particles (Alfa Aesar) characterized by a unimodal lognormal size distribution with a median of 250 nm were used as a magnetic filler for additional testing of the MI sensor prototype. The saturation magnetization of magnetite particles was 84 emu/g. Composites containing 5, 10, 30, 50, 70 (wt. %) were synthesized.

In experiments in-vitro, the FG biocompatibility was estimated by the proliferation rate of chondrocytes after 4 days of incubation. A primary culture of chondrocytes was obtained from the cartilage tissue of a sexually mature Wistar rat following the standard procedure¹¹ with the use of a Human Chondrocyte Growth Medium Kit (Cell applications, Inc, USA). 8 samples of PAAm gels and 8 samples of PAAm FG in the shape of disks (diameter ∼13 mm, 1mm thick) were placed into the wells of a polystyrene plate (Techno Plastic Products, Switzerland). A chondrocytes suspension in a growth medium was dispensed into all plate wells with a seeding density of 3000 viable cells/cm.² The plate with cells was incubated in a CO₂ incubator for 4 days, and the cell monolayer density was calculated after that. In the experiments with animals, the PAAm FG with 1.3 wt. % of MNPs concentration were examined. All animal experiments were performed in accordance with the International Guiding Principles for Biomedical Research Involving Animals (CIOMS, 1985) and approved by the Ethics Committee of the Ural State Medical University (Record #2 on Feb’28, 2020).

The multilayered elements for the biosensor were deposited by dc-sputtering with the application of an external magnetic field along the short side of the rectangular 10× 0.5 (mm²) element. FeNi/Ti and FeNi/Cu based structures were considered. The MI was measured in the “microstripe” line.¹⁰ The MI ratio for the impedance Z was calculated as follows: ΔZ=[Z(H)–Z (H_max)]/Z (H_max), H_max= 100 Oe. The magnetization curves were measured by a vibrating sample magnetometer. To analyze the MI effect in multilayers we proposed an electrodynamic model.¹² Since the multilayer length and width are much higher than its thickness, the field distribution within the layers can be found in the one-dimensional approximation by means of a solution of linearized Maxwell equations. To describe the field distribution outside the multilayer we use the approximate solution for the vector potential obtained previously.¹³ The MI response was found by taking into account the continuity conditions for the electric and magnetic fields at the interfaces between layers and at the multilayer surface.

The following cell counts after incubation were obtained for PAAm gel (control 1), PAAm + γ-Fe₂O₃ and polystyrene (control 2): 180 ± 60, 1500 ± 600 and 23500 ± 1800 cells/cm,² respectively. The data is presented in the form X ± m, where m is the standard error of the arithmetic average X, n = 72 (9 fields of view per well × 8 wells). The results obtained imply that the addition of MNPs to the network of PAAm gel results in a significant increase of the cell proliferation rate. Thus, the FG biocompatibility estimated by monolayer density of chondrocytes after 4 days of culturing is much higher than in the blank gel and closer to the tissue culture polystyrene. Qualitatively the same results were reported early with respect to human dermal fibroblasts.⁸ 

In vivo experiments were performed with 6 sexually mature male rabbits “Soviet Chinchilla” (certificate No. 2020/2KSh dated April 12, 2020), one of which was used as a donor for chondrocytes. During the surgery of the five animals, cylindrical osteochondral defects (5 mm in diameter and 5.5 mm in length) were formed on the anterior surface of the distal epimetaphysis femur of both limbs. Unsubstituted defects (4 defects) were the control for assessing the response of the surrounding tissues to surgery. PAAm-Fe₂O₃ (2 implants) and PAAm-Fe₂O₃ with allogeneic chondrocytes applied to the surface of the implant (4 implants) were tested. The cells were isolated from cartilage tissue and cultured using standard methods.¹⁴ The animals were taken out of the experiment 30 days after the day of operation. The defect area and adjacent bone and cartilage tissue of the host bone bed were examined by morphological and morphometric methods (Fig. 3). The biocompatibility was evaluated by the semi-quantitative assessment of the local biological effect of implants¹⁴ based on the determination of the presence of necrotic zones, counting the number of pro-inflammatory cells – polymorphonuclear leukocytes, lymphocytes, macrophages, plasma cells and giant multinucleated cells.

It was found (Fig. 4) that the introduction of implants (PAAm-Fe₂O₃ and PAAm-Fe₂O₃-allogeneic chondrocytes) caused the development of small necrotic foci in the adjacent tissues (in 3 out of 6 cases), activation of lymphocytes and macrophages. As expected, there was a tendency to an increase in the intensity of the local inflammatory reaction. An increase in the number of lymphocytes and macrophages in the maternal bed tissues was observed after the introduction of PAAm-Fe₂O₃-allogeneic chondrocytes in comparison with the control and the PAAm-Fe₂O₃. However, the severity of the signs of local inflammatory reaction did not significantly differ from control (p≤0.1). There were no signs of rejection or severe degradation of the implant during the introduction of both types of implants – there were no giant multinucleated cells, the number of neutrophils increased insignificantly compared to the control. Thus, under present conditions and taking into account the overall assessment of the selected criteria, the PAAm-Fe₂O₃ can be considered as a relatively inert¹⁵ or a biotolerant material.¹⁶ 

Coming back to the idea of combining positioning control and post operation monitoring for in vivo experiments with replacement of cartilage, we propose to develop an MI sensor for the detection of stray fields of MNPs of a FG implant. Fig. 5 shows an example of the characterization of multilayers, experimental frequency and field dependences of MI, including a FG detection example.

The MI effect in multilayers is governed by changes in the skin depth with the transverse permeability of FeNi layers. To find the transverse permeability we assume that FeNi layers have a uniaxial in-plane magnetic anisotropy, and the anisotropy field (H_a) and angle of deviation of the anisotropy axis from the transverse direction are the same for all FeNi layers. The transverse permeability of FeNi layers is obtained from a solution of the Landau–Lifshitz equation neglecting the contribution of the domain-walls motion. The developed model allows one to analyze the influence of multilayer parameters on the MI effect. Fig. 6(a) shows a calculated frequency dependence of the maximum MI ratio (ΔZ/Z)_max. The parameters of the FeNi layers used for calculations are: saturation magnetization M=750G, H_a=6Oe, anisotropy axis angle ψ=0.1π, conductivity σ=3×10¹⁶s⁻¹, and the Gilbert damping parameter κ=0.02. It follows from Fig. 6(a) that the value of (ΔZ/Z)_max decreases with the magnetic layer thickness. Furthermore, the MI effect increases when the conductivity of the spacers decreases (we assume that Cu conductivity σ_Cu=5×10¹⁷s⁻¹ and Ti conductivity σ_Ti=5×10¹⁶s⁻¹). The analysis demonstrated that the MI maximum can be achieved when the spacer conductivity becomes comparable in magnitude with the FeNi conductivity.¹⁷ It was found also that the further enhancement of the MI can be obtained in multilayers with thinner spacers and by using magnetic materials for spacers.¹⁸ The increase of (ΔZ/Z)_max in the multilayers with magnetic spacers is due to the fact that the difference between the skin depths in the spacers and in the FeNi layers is less pronounced: the field distribution within the multilayer becomes more uniform, leading to the MI increase (Fig.6). The presence of FG changes the field distribution within multilayered film structure and correspondingly affects the MI response. A model for the MI effect in multilayers with FG was developed recently.¹⁰ In accordance with the model, the FG layer has a double effect on the MI response. The first contribution to the MI effect is attributed to the high permittivity of the FG layer, which results in an increase of the MI in the multilayer with FG. The second contribution of the FG comes from stray fields of MNPs.

To describe qualitatively the effect of stray fields on the MI response, it is supposed that the FG layer induces a spatially homogeneous effective stray field H_p. The value of H_p is assumed to be proportional to the concentration of MNPs in the ferrogel.¹⁰ The field dependence of the MI ratio calculated at 150MHz for [Ti (6)/FeNi(100)]₅/Ti (6)/Cu(500)/Ti (6)/[FeNi(100)/Ti (6)]₅ structure (curve 1) and the multilayer with FG for different values of the effective stray field H_p (curves 2–5) is shown in Fig. 6(b). The frequency dependence of the maximum MI ratio (ΔZ/Z)_max for the same samples is shown in Fig. 6(c). In the presence of the gel without MNPs (H_p=0), the MI increases due to high permittivity of the gel. The effective stray field H_p increases with the concentration of magnetic nanoparticles in the FG, and the MI ratio decreases with an increase of H_p. The calculated dependences are in a qualitative agreement with experimental results (Fig. 5(d)). The stray fields change the magnetization distribution in the FeNi layers and affect the multilayer permeability allowing to measure the MNPs concentration in FG (Fig. 5(e)). As the next step to develop a magnetic biosensor for regenerative medicine purposes, we designed a scanning system to monitor the position and state of a FG implant at a distance. The first measurements of concentration dependence were made for the cylindrical samples of epoxy composites with commercial magnetite particles (Fig. 6(d)).

We described our experiments in the development of polyacrylamide ferrogels with γ-Fe₂O₃ magnetic nanoparticles produced by LTE. The examples provided in the field of regenerative medicine for cartilage replacement and magnetic biosensing using FeNi/Ti and FeNi/Cu based MI multilayers confirm the capacity of these functional materials for biomedical applications.

This work was supported by the RSF project 18-19-00090. We thank O.M. Samatov, B. Stadler, V.N. Lepalovskij, M.A. Korch, A.V. Svalov, D.G. Blisnets, A. Larrañaga, I. Orue, and M.N. Volochaev for support. Selected measurements were made at SGIKER UPV/EHU.

The data that support the findings of this study are available from the corresponding author upon reasonable request.