Fluorescence recovery after photobleaching (FRAP) is a method to measure molecular diffusion in tissues or gels, mainly for high molecular weight compounds. For this, the substance to be analyzed must be labeled with a fluorochrome. Mostly fluorescence-labeled proteins or FITC-dextranes (fluorescein isothiocyanate-dextranes) with different molecular sizes are used. The tissues or gels have to be soaked with this substance. This can be realized by storing the material in the fluorescence substance for some days or in case of a gel, to directly polymerize in the fluorescent substance. Then, a confocal laser is used to bleach out a certain area (mostly a line or a point) in the material. Because of diffusion, bleached molecules will move and change their position with fluorescent particles and the fluorescence recovers. After the bleaching process, the area will be scanned several times. The recovery time of fluorescence intensity is used to determine the diffusion coefficient of the substance in the material. For

Further examples for imaging methods for the determination of diffusion of molecules in skin are Fourier-transform-infrared (FTIR) spectroscopy [31, 32], two-photon fluorescence correlation spectroscopy in combination with fluorescence correlation spectroscopy (FCS) [33, 34] and optical coherence tomography [35]. These methods are noninvasive and nondestructive. Furthermore, some of them can detect molecules without fluorescence labeling. One big disadvantage is the equipment. For these imaging methods, special microscopes or also cost-

A method to investigate the penetration process of substances into the skin is tape stripping. After treatment of the skin with the substance of interest the stratum corneum is ripped of layer by layer with an adhesive film. Then, the amount of the substance can be analyzed. For this, there are different methods to determine the concentration of the substance. One method is to detect the substances directly on the film, for example titanium dioxide can be analyzed with X-ray fluorescent measurement and fluorescent-labeled substances can be detected via laser scanning microscopy. Another possibility is to remove the skin layer from the film and apply standardized analytical methods to determine the substance concentration. With tape stripping it is possible to observe where the substance of interest is localized and how deep

**permeation within the skin**

248 Biomaterials in Regenerative Medicine

The need to evaluate skin permeation, test cosmetic products and toxicologically screen topically applied compounds is evident. Historically, several millions of animal experiments have been performed worldwide to address this purpose [40]. Since animal experiments are under massive debate, ethical and regulatory issues, but also severe differences between animal and human data pushed the development and commercialization of diverse *in vitro* skin models [41, 42]. Human skin equivalents (HSEs) can be categorized into two main groups: the epidermis-only and full-thickness models. For both, the differentiation of keratinocytes and hence development of the various layers of the epidermis is important to model actual skin barrier properties more closely. In this context, the direct exposure to air as well as the culture media that supply nutrients for cell growth and differentiation from below has been found to be beneficial [43]. Growing cells on a porous membrane is one of the most commonly used ways to accomplish this air-liquid interface culture. According to the Organization for Economic Co-operation and Development (OECD) test guideline 431 (skin corrosion) and 439 (skin irritation), currently validated skin models include EpiSkin™ (L'oreal, France), EpiDerm™ SIT (MatTek Corporation, USA), SkinEthic™ RHE (SkinEthic laboratories, France), EpiCS® (CellSystems, Germany) and LabCyte EPI-MODEL24 SIT (Japan Tissue Engineering Co., Japan). These 3D skin models are all composed of one cell type only, the keratinocytes, mimicking the epidermis of native human skin and are especially advantageous with respect to high reproducibility [44]. However still not validated, there are also commercially available full thickness skin models composed of an additional dermal layer (e.g., GraftSkin®, EpiDermFT®, and Pheninon®). The models described are nowadays widely used for animalfree tests in drug development as well as, the chemical and cosmetics industries.

## **5. Determination of permeation and diffusion coefficients in membrane insert systems via measurement and simulation**

A common method for *in vitro* cultivation of skin models is the use of a membrane insert system. This system consists of a plastic vessel with a permeable membrane at the bottom. It enables the cultivation of skin tissue models in airlift on the membrane and guarantees the supply of nutrients from below. As the membrane insert system has two separate chambers, it can be used for permeation studies similar to the Franz diffusion cell. In the following, an experimental procedure for the determination of permeation coefficients in Transwell® systems (12 and 96 well) and simulations with COMSOL Multiphysics for estimation of diffusion coefficients will be discussed. Details can be found in [53].

structure enables the computation of different physical problems in one simulation. The permeation experiment was simulated with the module "transport diluted species", which uses the Fick's law to simulate diffusion processes. In order to determine the diffusion coefficient, a parameter optimization was performed with the "optimization module". Some simplifications have to be done in order to simulate the experiment. In the experiment, the permeating substance passes through a barrier and a membrane. For the simulation, these two phases were resumed as one homogenous material. It is not possible to resolve the different phases, as investigations that are more detailed would be necessary. The diffusion coefficient of the permeating substance in the liquid phases (in donor and acceptor) was

Measurement and Simulation of Permeation and Diffusion in Native and Cultivated Tissue…

represents the molecular distribution in the mixing process. Furthermore, the geometry of the membrane insert system was simplified. In reality, the system is slightly conical. The simplified geometry is a cylinder. All boundaries of the geometry were set as "no slip". The concave surface of the agarose gel was approximated with a spherical shape. In **Figure 3**,

Different membrane insert sizes influence the permeation within the system. Investigations with fluorescein sodium salt and 2% agarose gel in 96- and 12-well membrane insert systems are shown in **Figure 4a**. The time course of the acceptor concentration in the 12-well system was steeper in comparison to the 96-well system. Therefore, the fluorescein sodium salt permeates faster through the barrier in the 12-well system compared to the 96-well system. A reason for this is the concentration gradient in the gap between the membrane and the bottom of the receiver vessel. Because of the ratio "volume to permeation surface" and the gap size, the concentration will be balanced faster in case of the 12-well system. The gradient can be reduced by increasing the mixing frequency and amplitude. However, this is limited due to spillover of the liquid. By simulating these experiments, the different concentration distributions below the membrane can be visualized. In **Figure 4c** and **d** the concentration at different time points is plotted over the length below the membrane from the middle point to the edge of the receiver vessel (see red line on **Figure 4b**). The concentration difference between the middle and edge in 12-well systems is higher than that in 96-well systems. This indicates a better and faster concentration balance in the larger system, which explains the

The reproducibility of the suggested method is an important aspect. The permeation coefficient determined in different permeation experiments with fluorescein sodium salt through 2% agarose gel differed up to 40.9%. Although the value seems to be quite high, it is still within the range of deviation of permeability experiments with Franz diffusion cells reported in the literature [24]. It was found that small concentration variations during the preparation of the donor substance can cause large variation. By using one stock solution for different experiments, the deviation can be reduced to 29%. Further reasons for the deviation can also be small pipetting errors and variations of the barrier, for example, variation of the gel con-

/s. This parameter

251

http://dx.doi.org/10.5772/intechopen.72904

determined in preliminary mixing tests and was found to be 1 × 10−9 m2

**5.2. Influence of different settings and validation of the system.**

accelerated permeation.

centration or volume.

the geometry and the mesh of the 96- and 12-well Transwell® systems are shown.

#### **5.1. Measurement and simulation.**

The scheme of a permeation experiment in a membrane insert system is shown in **Figure 2**. The tissue barrier is established on the membrane. It is composed of agarose gel or 3D tissue to validate the method. The 3D tissues consist of a collagen matrix with human fibroblasts within and HaCaT cells on the top of the matrix. On top of the barrier, the donor is applied. The donor contains the substance to be analyzed, which permeates through the barrier. The acceptor, which collects the permeating substance, is located on the other side of the membrane in the receiver vessel. Temperature, humidity and mixing are parameters that influence the permeation and should be kept constant. It is recommended to perform the experiment on a shaker in an incubator with 37°C and ≥90% humidity (conditions for human cell culture). To avoid hydrodynamic pressure, the fluid surface in the insert (donor) and in the receiver vessel (acceptor) should be on the same level. The used volume for the experiment in 12 and 96 Transwell® systems is shown in **Table 1**. Sampling, like in the Franz diffusion cell, is difficult because of the small volume in the acceptor. A solution is the use of fluorescence-labeled substances, by which the permeate concentration can be detected via fluorescence measurement in the receiver vessel. A more elaborate possibility is to run several permeation experiments in parallel and take a sample from one vessel per time point. Then, the concentration of the substance can be measured analytically.

The permeation experiment was simulated with COMSOL Multiphysics. This program is based on the finite element method and offers different physic simulation modules. This

**Figure 2.** Schematic sketch of a membrane insert system.

**Table 1.** Liquid volume in the acceptor and the donor in different membrane insert systems with a barrier of 2 mm thickness.

structure enables the computation of different physical problems in one simulation. The permeation experiment was simulated with the module "transport diluted species", which uses the Fick's law to simulate diffusion processes. In order to determine the diffusion coefficient, a parameter optimization was performed with the "optimization module". Some simplifications have to be done in order to simulate the experiment. In the experiment, the permeating substance passes through a barrier and a membrane. For the simulation, these two phases were resumed as one homogenous material. It is not possible to resolve the different phases, as investigations that are more detailed would be necessary. The diffusion coefficient of the permeating substance in the liquid phases (in donor and acceptor) was determined in preliminary mixing tests and was found to be 1 × 10−9 m2 /s. This parameter represents the molecular distribution in the mixing process. Furthermore, the geometry of the membrane insert system was simplified. In reality, the system is slightly conical. The simplified geometry is a cylinder. All boundaries of the geometry were set as "no slip". The concave surface of the agarose gel was approximated with a spherical shape. In **Figure 3**, the geometry and the mesh of the 96- and 12-well Transwell® systems are shown.

#### **5.2. Influence of different settings and validation of the system.**

supply of nutrients from below. As the membrane insert system has two separate chambers, it can be used for permeation studies similar to the Franz diffusion cell. In the following, an experimental procedure for the determination of permeation coefficients in Transwell® systems (12 and 96 well) and simulations with COMSOL Multiphysics for estimation of diffusion

The scheme of a permeation experiment in a membrane insert system is shown in **Figure 2**. The tissue barrier is established on the membrane. It is composed of agarose gel or 3D tissue to validate the method. The 3D tissues consist of a collagen matrix with human fibroblasts within and HaCaT cells on the top of the matrix. On top of the barrier, the donor is applied. The donor contains the substance to be analyzed, which permeates through the barrier. The acceptor, which collects the permeating substance, is located on the other side of the membrane in the receiver vessel. Temperature, humidity and mixing are parameters that influence the permeation and should be kept constant. It is recommended to perform the experiment on a shaker in an incubator with 37°C and ≥90% humidity (conditions for human cell culture). To avoid hydrodynamic pressure, the fluid surface in the insert (donor) and in the receiver vessel (acceptor) should be on the same level. The used volume for the experiment in 12 and 96 Transwell® systems is shown in **Table 1**. Sampling, like in the Franz diffusion cell, is difficult because of the small volume in the acceptor. A solution is the use of fluorescence-labeled substances, by which the permeate concentration can be detected via fluorescence measurement in the receiver vessel. A more elaborate possibility is to run several permeation experiments in parallel and take a sample from one vessel per time point. Then, the concentration of the

The permeation experiment was simulated with COMSOL Multiphysics. This program is based on the finite element method and offers different physic simulation modules. This

**Table 1.** Liquid volume in the acceptor and the donor in different membrane insert systems with a barrier of 2 mm

coefficients will be discussed. Details can be found in [53].

**5.1. Measurement and simulation.**

250 Biomaterials in Regenerative Medicine

substance can be measured analytically.

**Figure 2.** Schematic sketch of a membrane insert system.

thickness.

Different membrane insert sizes influence the permeation within the system. Investigations with fluorescein sodium salt and 2% agarose gel in 96- and 12-well membrane insert systems are shown in **Figure 4a**. The time course of the acceptor concentration in the 12-well system was steeper in comparison to the 96-well system. Therefore, the fluorescein sodium salt permeates faster through the barrier in the 12-well system compared to the 96-well system. A reason for this is the concentration gradient in the gap between the membrane and the bottom of the receiver vessel. Because of the ratio "volume to permeation surface" and the gap size, the concentration will be balanced faster in case of the 12-well system. The gradient can be reduced by increasing the mixing frequency and amplitude. However, this is limited due to spillover of the liquid. By simulating these experiments, the different concentration distributions below the membrane can be visualized. In **Figure 4c** and **d** the concentration at different time points is plotted over the length below the membrane from the middle point to the edge of the receiver vessel (see red line on **Figure 4b**). The concentration difference between the middle and edge in 12-well systems is higher than that in 96-well systems. This indicates a better and faster concentration balance in the larger system, which explains the accelerated permeation.

The reproducibility of the suggested method is an important aspect. The permeation coefficient determined in different permeation experiments with fluorescein sodium salt through 2% agarose gel differed up to 40.9%. Although the value seems to be quite high, it is still within the range of deviation of permeability experiments with Franz diffusion cells reported in the literature [24]. It was found that small concentration variations during the preparation of the donor substance can cause large variation. By using one stock solution for different experiments, the deviation can be reduced to 29%. Further reasons for the deviation can also be small pipetting errors and variations of the barrier, for example, variation of the gel concentration or volume.

**Figure 3.** Geometry of the 96 a) and 12 b) Transwell® system and the used mesh c) and d) implemented in COMSOL **Mult**i**physics**.

a 12 Transwell® system. The 3D tissues consisted of a collagen matrix with different constellations of cell layers. Human primary fibroblasts were integrated into the collagen matrix and HaCaT cells were seeded on the top. A representative example of the 3D tissue is shown in

**Figure 4.** Permeation in 12- and 96-well membrane insert systems. (a) Experimental results of permeation of fluorescein sodium salt through 2% agarose. (b) Geometry of the 12 and 96 Transwell® system for the simulation and the position of the concentration measurement. (c) and (d) the concentration distribution at different time over the length at the red

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253

The results show that the permeation coefficient decreases when additional cell layers are added in the tissue model. The permeation coefficient of fluorescein sodium salt through collagen matrix (without cells) is 2.18 × 10−8 m/s, 1.85 × 10−8 m/s in tissue models with fibroblast and 1.67 × 10−8 m/s in models with HaCaT cells (see **Figure 6b)**. These results represent very

The particle size influences the permeation behavior through gels and biomaterials. It is well known that smaller molecules permeate faster through a matrix mesh than larger particles. This was already observed for permeation experiments through sclera [19], human epidermal membrane [45], human skin [33] and rat skin [3]. Fluorescein sodium salt and fluorescein isothiocyanate-dextrans (FITC-dextran) were used to vary the molecular size from 376 g/mol up

well the barrier function of keratinocytes of the skin [3].

lined position on (b) ((c) 96 and (d) 12 Transwell® system).

**Figure 5**.

The membrane itself also has an influence on the permeation. Same experiments as described above were carried out with different membranes in a 12 Transwell® system. Membranes consisting of polycarbonate (PC) with 0.4 and 3.0 μm pore size showed a quite similar permeation coefficient of 8.03 and 8.1 × 10−8 m/s. For polyethylene (PE) membranes with 0.4 μm pore size the mean value of the permeation coefficient was 5.94 × 10−8 m/s and for PE with 3.0 μm pore size it was 8.59 × 10−8 m/s (see **Figure 6a**). Except for the PE membrane with pore size of 3.0 μm there is no significant difference. The reason for this is the pore density of the membrane. In total, the pore surface of PE membranes with 0.4 μm pore size is 0.25 mm2 per 1 cm2 and for the other membranes 6.3–7.05 mm2 per 1 cm2 . The material of the membrane seems not to influence the permeation.

The suggested method is sensitive enough to determine a cover layer of HaCaT cells of 3D skin tissue. To prove this, in a permeation experiment different 3D tissue model was tested in

Measurement and Simulation of Permeation and Diffusion in Native and Cultivated Tissue… http://dx.doi.org/10.5772/intechopen.72904 253

**Figure 4.** Permeation in 12- and 96-well membrane insert systems. (a) Experimental results of permeation of fluorescein sodium salt through 2% agarose. (b) Geometry of the 12 and 96 Transwell® system for the simulation and the position of the concentration measurement. (c) and (d) the concentration distribution at different time over the length at the red lined position on (b) ((c) 96 and (d) 12 Transwell® system).

a 12 Transwell® system. The 3D tissues consisted of a collagen matrix with different constellations of cell layers. Human primary fibroblasts were integrated into the collagen matrix and HaCaT cells were seeded on the top. A representative example of the 3D tissue is shown in **Figure 5**.

The membrane itself also has an influence on the permeation. Same experiments as described above were carried out with different membranes in a 12 Transwell® system. Membranes consisting of polycarbonate (PC) with 0.4 and 3.0 μm pore size showed a quite similar permeation coefficient of 8.03 and 8.1 × 10−8 m/s. For polyethylene (PE) membranes with 0.4 μm pore size the mean value of the permeation coefficient was 5.94 × 10−8 m/s and for PE with 3.0 μm pore size it was 8.59 × 10−8 m/s (see **Figure 6a**). Except for the PE membrane with pore size of 3.0 μm there is no significant difference. The reason for this is the pore density of the membrane. In

**Figure 3.** Geometry of the 96 a) and 12 b) Transwell® system and the used mesh c) and d) implemented in COMSOL

The suggested method is sensitive enough to determine a cover layer of HaCaT cells of 3D skin tissue. To prove this, in a permeation experiment different 3D tissue model was tested in

per 1 cm2

. The material of the membrane seems not to

and for

total, the pore surface of PE membranes with 0.4 μm pore size is 0.25 mm2

per 1 cm2

the other membranes 6.3–7.05 mm2

influence the permeation.

**Mult**i**physics**.

252 Biomaterials in Regenerative Medicine

The results show that the permeation coefficient decreases when additional cell layers are added in the tissue model. The permeation coefficient of fluorescein sodium salt through collagen matrix (without cells) is 2.18 × 10−8 m/s, 1.85 × 10−8 m/s in tissue models with fibroblast and 1.67 × 10−8 m/s in models with HaCaT cells (see **Figure 6b)**. These results represent very well the barrier function of keratinocytes of the skin [3].

The particle size influences the permeation behavior through gels and biomaterials. It is well known that smaller molecules permeate faster through a matrix mesh than larger particles. This was already observed for permeation experiments through sclera [19], human epidermal membrane [45], human skin [33] and rat skin [3]. Fluorescein sodium salt and fluorescein isothiocyanate-dextrans (FITC-dextran) were used to vary the molecular size from 376 g/mol up

in **Figure 8a**. In the case of larger molecular size, the simulation is different from experimental results. A closer look shows that the experimental data increased and flattened faster compared to the simulation (see **Figure 8b)**. A possible reason for this effect could be the presence of particle size distribution in the substance. The migration of smaller particles reduces the lag time in the beginning of the permeation process, where large particles can increase the friction and slows down the diffusion. Especially the second effect leads to abnormal diffusion [46, 47].

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255

**Figure 7.** Permeability coefficient plotted over stokes radius. Results of permeation experiment in 96 Transwell® system

**Figure 8.** Simulation of permeation experiment with different molecular sizes. Results of permeation experiments and simulations with COMSOL **Multiphysics** of (a) fluorescein sodium salt and (b) FITC dextran 10,000 mg/mol in 96

This cannot be simulated with equations based on Fick's law.

with fluorescein sodium salt and FITC-dextran through 2% agarose gel.

Transwell systems through 2% agarose gel.

**Figure 5.** Hematoxylin and eosin stain of 3D tissue model consisting of collagen matrix with fibroblast and HaCaT seeded on the top.

**Figure 6.** Permeation through different membranes and 3D skin tissues. (a) Results of permeation experiment with fluorescein sodium salt through 2% agarose gel on different membranes (n = 3). The 12 Transwell® system was used and membranes consisting of polycarbonate (PC) and polyethylene (PE) with pore sizes of 3.0 and 0.4 μm were tested. (b) Results of permeation experiments with fluorescein sodium salt through different 3D tissues in a 12 Transwell® system (n = 6). The 3D skin tissue consisted of collagen (Col.), collagen with primary human fibroblast (Col. + F.), collagen with HaCaTs (Col. + H.) and collagen with primary human fibroblast and HaCaTs (Col. F. + H.).

to 40,000 g/mol for permeation experiments in a 96 Transwell® system. The results show similar correlations between permeation coefficient and molecular size as the studies mentioned above. There is almost a linear relationship between these two factors, which is well described by the Navier–Stokes equation (see **Figure 7**). An exception is FITC-dextran 40,000 g/mol, which deviates from the linearity.

These experiments were simulated with COMSOL Multiphysics, where the diffusion coefficient is fitted on the experimental data. The simulation based on Fick's law is quite accurate for the permeation of substances with a small molecular size from 376 g/mol up to 4000 g/ mol. The simulation shows good agreement with the experiment which is exemplary shown in **Figure 8a**. In the case of larger molecular size, the simulation is different from experimental results. A closer look shows that the experimental data increased and flattened faster compared to the simulation (see **Figure 8b)**. A possible reason for this effect could be the presence of particle size distribution in the substance. The migration of smaller particles reduces the lag time in the beginning of the permeation process, where large particles can increase the friction and slows down the diffusion. Especially the second effect leads to abnormal diffusion [46, 47]. This cannot be simulated with equations based on Fick's law.

**Figure 7.** Permeability coefficient plotted over stokes radius. Results of permeation experiment in 96 Transwell® system with fluorescein sodium salt and FITC-dextran through 2% agarose gel.

to 40,000 g/mol for permeation experiments in a 96 Transwell® system. The results show similar correlations between permeation coefficient and molecular size as the studies mentioned above. There is almost a linear relationship between these two factors, which is well described by the Navier–Stokes equation (see **Figure 7**). An exception is FITC-dextran 40,000 g/mol,

HaCaTs (Col. + H.) and collagen with primary human fibroblast and HaCaTs (Col. F. + H.).

**Figure 6.** Permeation through different membranes and 3D skin tissues. (a) Results of permeation experiment with fluorescein sodium salt through 2% agarose gel on different membranes (n = 3). The 12 Transwell® system was used and membranes consisting of polycarbonate (PC) and polyethylene (PE) with pore sizes of 3.0 and 0.4 μm were tested. (b) Results of permeation experiments with fluorescein sodium salt through different 3D tissues in a 12 Transwell® system (n = 6). The 3D skin tissue consisted of collagen (Col.), collagen with primary human fibroblast (Col. + F.), collagen with

**Figure 5.** Hematoxylin and eosin stain of 3D tissue model consisting of collagen matrix with fibroblast and HaCaT

These experiments were simulated with COMSOL Multiphysics, where the diffusion coefficient is fitted on the experimental data. The simulation based on Fick's law is quite accurate for the permeation of substances with a small molecular size from 376 g/mol up to 4000 g/ mol. The simulation shows good agreement with the experiment which is exemplary shown

which deviates from the linearity.

seeded on the top.

254 Biomaterials in Regenerative Medicine

**Figure 8.** Simulation of permeation experiment with different molecular sizes. Results of permeation experiments and simulations with COMSOL **Multiphysics** of (a) fluorescein sodium salt and (b) FITC dextran 10,000 mg/mol in 96 Transwell systems through 2% agarose gel.

## **6. Conclusion and prospective**

The studies have shown that the membrane insert system is a possible alternative for permeation studies. An advantage of the system is the small size. The membrane insert system of 96-well plates from CORNING has a cultivation surface of 0.143 cm<sup>2</sup> and a height of 2 cm. This reduces significantly the number of cells, materials and substances needed for the cultivation. In comparison to the Franz diffusion cell, the handling of such a system is easier and one experiment can be run with a large specimen number. A time-intensive mounting process of the samples (skin) is no longer required and the experiment can directly execute in the system. The sensitivity of this system is good enough to differ between 3D tissues as well as different cell layers and to detect different molecular sizes of the substance.

FRAP fluorescence recovery after photobleaching

HaCaT human adult low calcium high temperature

, Katharina Schimek<sup>2</sup>

Physiology. 2010;**23**(2):105-112. DOI: 10.1159/000265681

Release. 1987;**6**(1):85-97. DOI: 10.1016/0168-3659(87)90066-6

1-18. DOI: 10.1016/S0378-5173(01)00731-1

2194. DOI: 10.1002/jps.20695

**208**:7-11. PubMed ID: 10884933

ID: 11509908

\*Address all correspondence to<EMAIL_ADDRESS>
OECD organization for economic co-operation and development

, Uwe Marx2

1 Institute of Bioprocess and Biosystems Engineering, Hamburg University of Technology,

[1] Ackermann K, Lombardi Borgia S, Korting H, Mewes K, Schäfer-Korting M.The Phenion® full-thickness skin model for percutaneous absorption testing. Skin Pharmacology and

[2] Asbill C. Evaluation of a human bio-engineered skin equivalent for drug permeation studies. Pharmaceutical Research. 2000;**17**(9):1092-1097. DOI: 10.1023/A:1026405712870

[3] Hadgraft J. Skin, the final frontier. International Journal of Pharmaceutics. 2001;**224**(1-2):

[4] Barry BW. Mode of action of penetration enhancers in human skin. Journal of Controlled

[5] Valenzuela P, Simon JA. Nanoparticle delivery for transdermal HRT. Nanomedicine: Nanotechnology, Biology and Medicine. 2012;**8**:83-89. DOI: 10.1016/j.nano.2012.05.008

[6] Barbero AM, Frasch HF. Transcellular route of diffusion through stratum corneum: Results from finite element models. Journal of Pharmaceutical Sciences. 2006;**95**(10):2186-

[7] Bouwstra J, Pilgram G, Gooris G, Koerten H, Ponec M. New aspects of the skin barrier organization. Skin Pharmacology and Applied Skin Physiology. 2001;**14**:52-62. PubMed

[8] Wertz PW. Lipids and barrier function of the skin. Acta Dermato-Venereologica. 2000;

and Ralf Pörtner1

Measurement and Simulation of Permeation and Diffusion in Native and Cultivated Tissue…

\*

http://dx.doi.org/10.5772/intechopen.72904

257

FTIR Fourier-transform-infrared

HSE human skin equivalents

**Author details**

Hao-Hsiang Hsu1

Hamburg, Germany

**References**

2 TissUse GmbH, Berlin, Germany

It should be considered that the permeation is detected through a membrane and the size of the system influences the permeation properties. The specimen has to cover up the whole membrane, otherwise the substance will pass by. Unlike the imaging and stripping method, it is not possible to measure the diffusion and penetration inside the membrane insert system. Alternatively, the diffusion can be calculated or estimated by simulation. Furthermore, this method can be used to investigate changes in the permeation behavior of the skin model during the cultivation or it can also be adapted for other systems, which use membrane insert systems. An example is the Two-Organ-on-a-Chip, a variant of TissUse's Multi-Organ-Chip platform [48–52]. This device enables the integration of skin models in a membrane insert system. Therefore this method can be used to investigate the permeation process into an organon-a-chip system in order to understand the substance distribution.

With the help of the simulation in COMSOL Multiphysics, it is possible to calculate the diffusion process in the membrane insert system. It is limited to small particle sizes and normal diffusion described by Fick's law. Otherwise, it is possible to optimize the simulation by integration of abnormal diffusion. Furthermore, the simulation is an attractive tool to support the experiments. On the one hand, it can be used to understand physical phenomena and to reduce experimental effort. On the other hand, it is modular and can be integrated into a more complex system to support permeation studies.

## **Acknowledgements**

This work was created with financial support from Deutsche Forschungsgemeinschaft (DFG) under grant No. PO413/12-1 and LA 1028/7-1.

## **Appendices and nomenclatures**

EC European regulation

FITC dextran fluorescein isothiocyanate-dextrans

## **Author details**

**6. Conclusion and prospective**

256 Biomaterials in Regenerative Medicine

The studies have shown that the membrane insert system is a possible alternative for permeation studies. An advantage of the system is the small size. The membrane insert system

This reduces significantly the number of cells, materials and substances needed for the cultivation. In comparison to the Franz diffusion cell, the handling of such a system is easier and one experiment can be run with a large specimen number. A time-intensive mounting process of the samples (skin) is no longer required and the experiment can directly execute in the system. The sensitivity of this system is good enough to differ between 3D tissues as well as

It should be considered that the permeation is detected through a membrane and the size of the system influences the permeation properties. The specimen has to cover up the whole membrane, otherwise the substance will pass by. Unlike the imaging and stripping method, it is not possible to measure the diffusion and penetration inside the membrane insert system. Alternatively, the diffusion can be calculated or estimated by simulation. Furthermore, this method can be used to investigate changes in the permeation behavior of the skin model during the cultivation or it can also be adapted for other systems, which use membrane insert systems. An example is the Two-Organ-on-a-Chip, a variant of TissUse's Multi-Organ-Chip platform [48–52]. This device enables the integration of skin models in a membrane insert system. Therefore this method can be used to investigate the permeation process into an organ-

With the help of the simulation in COMSOL Multiphysics, it is possible to calculate the diffusion process in the membrane insert system. It is limited to small particle sizes and normal diffusion described by Fick's law. Otherwise, it is possible to optimize the simulation by integration of abnormal diffusion. Furthermore, the simulation is an attractive tool to support the experiments. On the one hand, it can be used to understand physical phenomena and to reduce experimental effort. On the other hand, it is modular and can be integrated into a more

This work was created with financial support from Deutsche Forschungsgemeinschaft (DFG)

and a height of 2 cm.

of 96-well plates from CORNING has a cultivation surface of 0.143 cm<sup>2</sup>

different cell layers and to detect different molecular sizes of the substance.

on-a-chip system in order to understand the substance distribution.

complex system to support permeation studies.

under grant No. PO413/12-1 and LA 1028/7-1.

**Appendices and nomenclatures**

EC European regulation

FITC dextran fluorescein isothiocyanate-dextrans

**Acknowledgements**

Hao-Hsiang Hsu1 , Katharina Schimek<sup>2</sup> , Uwe Marx2 and Ralf Pörtner1 \*

\*Address all correspondence to<EMAIL_ADDRESS>
1 Institute of Bioprocess and Biosystems Engineering, Hamburg University of Technology, Hamburg, Germany

2 TissUse GmbH, Berlin, Germany

## **References**

[9] Tsukita S, Furuse M, Itoh M. Multifunctional strands in tight junctions. Nature Reviews. Molecular Cell Biology. 2001;**2**(4):285-293. DOI: 10.1038/35067088

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260 Biomaterials in Regenerative Medicine

**Chapter 11**

**Provisional chapter**

**Biomaterials for Tendon/Ligament and Skin**

**Biomaterials for Tendon/Ligament and Skin**

DOI: 10.5772/intechopen.69716

Tendon/ligament injury or skin injuries due to diseases, trauma, and surgery are common. Timely functional repair and tissue regeneration is a key to improve the quality of life of the patient while reducing health care cost. Tendon/ligament/skin is also enriched in a common extracellular matrix (ECM), collagen I, III, and elastin. Tissue engineering and regenerative medicine, the combination of (stem) cells, growth factors, and biomaterial scaffolds, is an emergent field, which has attracted substantial attention over the years. Biomaterials are considered the foundation of regenerative medicine. A key to find a new solution to tendon/ligament/skin healing is to synthesize new functional biomaterials, which have better biomechanical properties, biodegradability, and cell supporting prop‐ erties. This chapter will review existing FDA‐approved biomaterial‐based therapy, as well

> © 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

From material science point of view, tendon/ligament and skin tissue are similar in that they are mainly composed of collagen and elastin. Up to 80% of the dry weight is collagen type I. However, their micro and hierarchical structures and functions are very different from each other. Tendon/ligament is composed of densely packed, aligned collagen fiber bundles, whereas skin is composed of a layered structure of collagen random naonfibril network. **Table 1**

The regeneration of a large size, lost tendon/ligament, and skin tissue normally involves a type of stem cells/progenitor cells of endogenous origin or exogenous origin combined with a

summarizes the key properties of tendon/ligament and skin tissue in human.

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69716

as those in development.

**Keywords:** biomaterials, tendon, skin, regeneration

**Regeneration**

**Regeneration**

Xingguo Cheng

**Abstract**

**1. Introduction**

Xingguo Cheng

**Provisional chapter**

## **Biomaterials for Tendon/Ligament and Skin Regeneration Regeneration**

**Biomaterials for Tendon/Ligament and Skin**

DOI: 10.5772/intechopen.69716

#### Xingguo Cheng Xingguo Cheng Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69716

#### **Abstract**

Tendon/ligament injury or skin injuries due to diseases, trauma, and surgery are common. Timely functional repair and tissue regeneration is a key to improve the quality of life of the patient while reducing health care cost. Tendon/ligament/skin is also enriched in a common extracellular matrix (ECM), collagen I, III, and elastin. Tissue engineering and regenerative medicine, the combination of (stem) cells, growth factors, and biomaterial scaffolds, is an emergent field, which has attracted substantial attention over the years. Biomaterials are considered the foundation of regenerative medicine. A key to find a new solution to tendon/ligament/skin healing is to synthesize new functional biomaterials, which have better biomechanical properties, biodegradability, and cell supporting prop‐ erties. This chapter will review existing FDA‐approved biomaterial‐based therapy, as well as those in development.

**Keywords:** biomaterials, tendon, skin, regeneration

## **1. Introduction**

From material science point of view, tendon/ligament and skin tissue are similar in that they are mainly composed of collagen and elastin. Up to 80% of the dry weight is collagen type I. However, their micro and hierarchical structures and functions are very different from each other. Tendon/ligament is composed of densely packed, aligned collagen fiber bundles, whereas skin is composed of a layered structure of collagen random naonfibril network. **Table 1** summarizes the key properties of tendon/ligament and skin tissue in human.

The regeneration of a large size, lost tendon/ligament, and skin tissue normally involves a type of stem cells/progenitor cells of endogenous origin or exogenous origin combined with a

effect, whereas pure PRP (P‐PRP) without leucocytes have anabolic effects and results in over‐ proliferation and scar tissue formation. The complex interaction between PRP and stem cells may explain the different preclinical outcome and warrant a large random clinical trial.

Biomaterials for Tendon/Ligament and Skin Regeneration

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265

Fibrin can be derived from human plasma. After addition of thrombin, it will form a gel, a process used for blood clotting. Stem cells can be added together with fibrin and thrombin and sprayed onto the wound for the promotion of wound closure and healing. The fibrin combined with bone marrow stem cells spray was successfully tested to prevent ulceration and acceler‐ ate wound closure in mice [4]. A small human clinical trial showed this approach accelerates wound closure and resurfacing without adverse effects [5]. Poly(ethylene glycol) PEG‐modified fibrin combined with adipose‐derived stem cells (ADSCs) also showed promising results in

Human placenta‐derived biomaterial is unique in that it has immune privilege while it con‐ tains multiple growth factors. Human ADSCs (hADSCs) seeded onto radiosterilized human amnion are viable and can proliferate. These cells are able to migrate over these scaffolds as demonstrated by using time‐lapse microscopy. In addition, the scaffolds induce hADSCs to secrete interleukin‐10, an important negative regulator of inflammation [7]. This suggests that placenta‐derived biomaterial may be a good substrate for stem cells and used for skin/tendon

3D micronized (300–600 μm) amniotic membrane (mAM) was made by means of repeated freeze‐thawing cycles to deplete cell components and homogenized with a macrohomoge‐ nizer in liquid nitrogen. These mAM loaded with epidermal stem cells (ESCs) (ESC‐mAM) was further transplanted to full‐thickness skin defects in nude mice. ESCs survived well and formed a new epidermis. Four weeks after transplantation, papilla‐like structures were observed, and collagen fibers were well and regularly arranged in the newly formed dermal layer. In conclusion, the mAM as a novel natural microcarrier possesses an intact basement

A recent study showed that decellularized matrix from different tissues (tendon, bone, and skin) affect the differentiation of stem cell in a different way. Decellularized bone matrix may induce the undesirable osteogenic differentiation of stem cells, while tendon or skin matrix does not have such an effect [9]. The ECM components provide a niche for proper differentia‐ tion of stem cells. For example, ECM without biglycan (Bgn) and fibromodulin (Fmod) will affect the differentiation of tendon stem cells by modulating bone morphogenetic protein signaling and impairs tendon formation *in vivo* [10]. Autologous origin, but decelluarized der‐ mal matrix using trypsin and Trion X‐100, after combined with ADSCs, was found to enhance

**2.2. FDA‐approved ECM grafts for tendon augmentation and skin regeneration**

wound healing in a murine model of a full‐thickness skin defect [11].

*2.1.2.  Amniotic membrane*

membrane structure and bioactivities [8].

**Table 1.** A general comparison of two different connective tissues: tendon/ligament and skin.

biomaterial. Stem cells can be derived from tendon/ligament or skin, or from embryo, pla‐ centa, adipose tissue, bone marrow, umbilical cord, etc. This review focuses only on regenera‐ tion using exogenous stem cells coupled with a biomaterial matrix/carrier implanted to the wound site. Biomaterial can play a key role in protection of the cells from dehydration while it serves as a temporary substrate for stem cells to proliferate, or differentiate, and synthesize tissue‐specific matrix. The morphology, topography, composition, stiffness of biomaterials may play a key role in controlling the differentiation of stem cells, in addition to biochemical factors, mechanical cues, or genetic/cellular cues. This chapter focuses on biomaterial explored for tendon/ligament/skin tissue regeneration applications.

#### **1.1. Research methods**

We performed a comprehensive search of PubMed using keywords "tendon," "ligament," "skin," "regeneration," "scaffold", over the years 1970–2016. All articles relevant to the subject were retrieved, and their bibliographies hand searched for further references in the context to biomaterials for tendon/ligament/skin regeneration

## **2. Results**

#### **2.1. Biomaterial directly derived from patients**

#### *2.1.1. PRP*

Platelet‐rich plasma (PRP) is derived from blood and PRP gel is widely used for tendon/ ligament repair. Recently, PRP was combined with adipose‐derived stem cells (ADSCs) and it was found that PRP combined with stem cells resulted in improved mechanical strength in a rabbit tendon model compared to PRP gel alone [1]. Similar results were also observed using PRP with tendon‐derived stem cells (TDSCs) [2]. However, in a sheep model, no differ‐ ences were observed between the PRP group and PRP‐stem cell group [3]. This approach is highly translational, since both autologous stem cells and PRP can be obtained from the same patients. The concern may be the leucocytes‐containing PRP (L‐PRP) that have a catabolic effect, whereas pure PRP (P‐PRP) without leucocytes have anabolic effects and results in over‐ proliferation and scar tissue formation. The complex interaction between PRP and stem cells may explain the different preclinical outcome and warrant a large random clinical trial.

#### *2.1.2. Fibrin*

biomaterial. Stem cells can be derived from tendon/ligament or skin, or from embryo, pla‐ centa, adipose tissue, bone marrow, umbilical cord, etc. This review focuses only on regenera‐ tion using exogenous stem cells coupled with a biomaterial matrix/carrier implanted to the wound site. Biomaterial can play a key role in protection of the cells from dehydration while it serves as a temporary substrate for stem cells to proliferate, or differentiate, and synthesize tissue‐specific matrix. The morphology, topography, composition, stiffness of biomaterials may play a key role in controlling the differentiation of stem cells, in addition to biochemical factors, mechanical cues, or genetic/cellular cues. This chapter focuses on biomaterial explored

**Tendon/ligament Skin**

Structure 1D Aligned fiber bundles (nano‐macro) 2D random nanofibrils (weave basket

pattern)

Main composition >80% Collagen, ∼5% elastin elastin >80% Collagen, ∼5% elastin Minor composition Fibronectin, proteoglycan Fibronectin, proteoglycan Main cells Fibroblast Fibroblast and keratinocytes

Vasculature Few Abundant

Mechanical properties Strong and elastic Weak and elastic

**Table 1.** A general comparison of two different connective tissues: tendon/ligament and skin.

We performed a comprehensive search of PubMed using keywords "tendon," "ligament," "skin," "regeneration," "scaffold", over the years 1970–2016. All articles relevant to the subject were retrieved, and their bibliographies hand searched for further references in the context to

Platelet‐rich plasma (PRP) is derived from blood and PRP gel is widely used for tendon/ ligament repair. Recently, PRP was combined with adipose‐derived stem cells (ADSCs) and it was found that PRP combined with stem cells resulted in improved mechanical strength in a rabbit tendon model compared to PRP gel alone [1]. Similar results were also observed using PRP with tendon‐derived stem cells (TDSCs) [2]. However, in a sheep model, no differ‐ ences were observed between the PRP group and PRP‐stem cell group [3]. This approach is highly translational, since both autologous stem cells and PRP can be obtained from the same patients. The concern may be the leucocytes‐containing PRP (L‐PRP) that have a catabolic

for tendon/ligament/skin tissue regeneration applications.

biomaterials for tendon/ligament/skin regeneration

**2.1. Biomaterial directly derived from patients**

**1.1. Research methods**

264 Biomaterials in Regenerative Medicine

**2. Results**

*2.1.1. PRP*

Fibrin can be derived from human plasma. After addition of thrombin, it will form a gel, a process used for blood clotting. Stem cells can be added together with fibrin and thrombin and sprayed onto the wound for the promotion of wound closure and healing. The fibrin combined with bone marrow stem cells spray was successfully tested to prevent ulceration and acceler‐ ate wound closure in mice [4]. A small human clinical trial showed this approach accelerates wound closure and resurfacing without adverse effects [5]. Poly(ethylene glycol) PEG‐modified fibrin combined with adipose‐derived stem cells (ADSCs) also showed promising results in a pig burn model [6].

#### *2.1.3. Amniotic membrane*

Human placenta‐derived biomaterial is unique in that it has immune privilege while it con‐ tains multiple growth factors. Human ADSCs (hADSCs) seeded onto radiosterilized human amnion are viable and can proliferate. These cells are able to migrate over these scaffolds as demonstrated by using time‐lapse microscopy. In addition, the scaffolds induce hADSCs to secrete interleukin‐10, an important negative regulator of inflammation [7]. This suggests that placenta‐derived biomaterial may be a good substrate for stem cells and used for skin/tendon applications.

3D micronized (300–600 μm) amniotic membrane (mAM) was made by means of repeated freeze‐thawing cycles to deplete cell components and homogenized with a macrohomoge‐ nizer in liquid nitrogen. These mAM loaded with epidermal stem cells (ESCs) (ESC‐mAM) was further transplanted to full‐thickness skin defects in nude mice. ESCs survived well and formed a new epidermis. Four weeks after transplantation, papilla‐like structures were observed, and collagen fibers were well and regularly arranged in the newly formed dermal layer. In conclusion, the mAM as a novel natural microcarrier possesses an intact basement membrane structure and bioactivities [8].

#### **2.2. FDA‐approved ECM grafts for tendon augmentation and skin regeneration**

A recent study showed that decellularized matrix from different tissues (tendon, bone, and skin) affect the differentiation of stem cell in a different way. Decellularized bone matrix may induce the undesirable osteogenic differentiation of stem cells, while tendon or skin matrix does not have such an effect [9]. The ECM components provide a niche for proper differentia‐ tion of stem cells. For example, ECM without biglycan (Bgn) and fibromodulin (Fmod) will affect the differentiation of tendon stem cells by modulating bone morphogenetic protein signaling and impairs tendon formation *in vivo* [10]. Autologous origin, but decelluarized der‐ mal matrix using trypsin and Trion X‐100, after combined with ADSCs, was found to enhance wound healing in a murine model of a full‐thickness skin defect [11].

Many FDA‐approved human or animal decelluarized tissue matrixes have been approved for direct tendon/ligament/skin repair without any (stem) cells (**Table 2**). The key advantages of decellularized tissue grafts are that they largely maintain the main architecture, composi‐ tion, and mechanical properties of native tissues. These allografts/xenografts are processed to remove immunogenic cells, DNAs, and certain immunogenic molecules. Typical problems are that these grafts are slower to repair the tissue and some fail to restore the proper func‐ tions (e.g., scarring). For tissue regeneration using these ECM biological grafts, stem cells may need to be reseeded onto the grafts for recellularization. For a dense tendon allograft, direct cell seeding may be difficult. The recellularization onto the surface may be achieved using a cell‐loaded gel coating [12]. However, it is highly desirable to get cells inside the grafts as well. Thus, ECM grafts may be processed to have a much higher porosity than the original tissue. Proration/incision into the ECM grafts may help the penetration of cells and nutrients. Instead of being coated with cell‐laden gel, an interesting approach is to use a stem cell‐sheet to wrap around a frozen tendon graft for implantation [13]. Interesting, a case report showed that a dermal allograft combined with PRP and autologous mesenchymal stem cells (MSCs) derived from peripheral blood (PB‐MSCs) resulted in enhanced healing of human rotator cuff [14].

patch (SIS) as carrier of ADSCs significantly increases the survival of stem cells as compared to direct injection of ADSCs into the skin wounds [18]. There may be a synergistic angiogen‐ esis promoting effects of biologic graft with stem cells, which may be important for tendon/

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Another interesting approach is to use ECM directly secreted by the cells as a carrier for stem cells. For example, stem cells can be incubated at 37°C in a temperature‐responsive flask (e.g., upcellTM, Cellseed, Japan) and they will product ECM after addition of ascorbic acid. The cell sheets can be lifted at room temperature since the Poly‐n‐isopropylacrylamide (Poly‐NIPAM)‐ based substrate will become soluble at lower temperature. The cell‐ECM sheet has been explored for promotion of tendon/ligament healing [20] as well as diabetic skin wound healing [21].

(5 mm × 2 mm) collagen sponge scaffolds (Zimmer Dental) were used for the culture of (BM) MSCs. Cell‐seeded scaffolds were placed in culture dishes and incubated for 2 hours in a minimum volume of growth medium, after which more medium was applied to submerge the scaffolds. After an additional 24‐hour culture, cells seeded in scaffolds were treated with 10 ng/mL of recombinant (BMP) 12 for 12 hours. The medium was then replaced with fresh growth medium and scaffolds were either cultured for an additional 7 days or immediately implanted into partial calcaneal tendon defects in rats. It was shown that after 21 days, the BMP12‐treated, collagen‐cell scaffold results in robust formation of tendon‐like tissue [22]. Similarly, a collagen carrier combined with ASDCs showed they did not improve the biome‐ chanical properties of the tendon‐to‐bone healing. However, the ADSCs group showed less inflammation, which may lead to a more elastic repair and less scarred healing in a rat model.

Integra® bilayer wound Matrix (LifeSciences Corp., Plainsboro, New Jersey) is a dermal acellular analog composed of bovine collagen type I crosslinked with glycosaminoglycans. Importantly, inclusion of WJ‐MSC into Integra induced significant up‐regulation of prototypical angiogenic and healing factors, stimulating pleiotropic aspects of neovascularization in experimental set‐ tings of angiogenesis, without altering the inflammatory response in the animals, thus dem‐ onstrating their potential benefit in therapeutic care of wounds and skin grafts [23]. Recently, Integra dermal matrix scaffold engineered with adult mesenchymal stem cells and platelet‐rich

plasma was investigated *in vitro* and demonstrated promising preliminary results [24].

Biopolymer matrix such as collagen, gelatin, hyaluronic acid, chitosan, silk, cellulose are the essential ECM components of human, animal, or plant cells. They have showed great promise for attachment, proliferation, and differentiation of stem cells and have been explored for

skin wound healing [19].

**2.3. FDA‐approved biomaterial**

*2.3.1. Collagen sponge scaffold*

*2.3.2. Integra bilayer*

**2.4. Synthetic biopolymer matrix**

tendon and skin regeneration.

Some of the FDA‐approved ECM biomatrices were combined with stem cells, and investigated for tendon/ligament/skin regeneration applications. Human acellular allograft (Alloderm) was investigated for direct cell seeding using ADSCs. It was concluded that human ADSCs can attach to Alloderm with the dermis side up in a petri dish [15]. ADSC seeded onto Alloderm was also implanted *in vivo* for skin regenerations with promising results [16]. Strattice was evaluated for seeding with rat MSCs [17]. Thus, the stem cell‐seeded biologic graft can be used as a tendon wrap or a skin regeneration material. A study was performed to compare the survival and proliferation of stem cells via bioluminescent imaging. The use of biologic graft

**Table 2.** An example list of biologic ECM grafts with potential for tendon/skin regeneration.

patch (SIS) as carrier of ADSCs significantly increases the survival of stem cells as compared to direct injection of ADSCs into the skin wounds [18]. There may be a synergistic angiogen‐ esis promoting effects of biologic graft with stem cells, which may be important for tendon/ skin wound healing [19].

Another interesting approach is to use ECM directly secreted by the cells as a carrier for stem cells. For example, stem cells can be incubated at 37°C in a temperature‐responsive flask (e.g., upcellTM, Cellseed, Japan) and they will product ECM after addition of ascorbic acid. The cell sheets can be lifted at room temperature since the Poly‐n‐isopropylacrylamide (Poly‐NIPAM)‐ based substrate will become soluble at lower temperature. The cell‐ECM sheet has been explored for promotion of tendon/ligament healing [20] as well as diabetic skin wound healing [21].

#### **2.3. FDA‐approved biomaterial**

#### *2.3.1. Collagen sponge scaffold*

Many FDA‐approved human or animal decelluarized tissue matrixes have been approved for direct tendon/ligament/skin repair without any (stem) cells (**Table 2**). The key advantages of decellularized tissue grafts are that they largely maintain the main architecture, composi‐ tion, and mechanical properties of native tissues. These allografts/xenografts are processed to remove immunogenic cells, DNAs, and certain immunogenic molecules. Typical problems are that these grafts are slower to repair the tissue and some fail to restore the proper func‐ tions (e.g., scarring). For tissue regeneration using these ECM biological grafts, stem cells may need to be reseeded onto the grafts for recellularization. For a dense tendon allograft, direct cell seeding may be difficult. The recellularization onto the surface may be achieved using a cell‐loaded gel coating [12]. However, it is highly desirable to get cells inside the grafts as well. Thus, ECM grafts may be processed to have a much higher porosity than the original tissue. Proration/incision into the ECM grafts may help the penetration of cells and nutrients. Instead of being coated with cell‐laden gel, an interesting approach is to use a stem cell‐sheet to wrap around a frozen tendon graft for implantation [13]. Interesting, a case report showed that a dermal allograft combined with PRP and autologous mesenchymal stem cells (MSCs) derived from peripheral blood (PB‐MSCs) resulted in enhanced healing of

Some of the FDA‐approved ECM biomatrices were combined with stem cells, and investigated for tendon/ligament/skin regeneration applications. Human acellular allograft (Alloderm) was investigated for direct cell seeding using ADSCs. It was concluded that human ADSCs can attach to Alloderm with the dermis side up in a petri dish [15]. ADSC seeded onto Alloderm was also implanted *in vivo* for skin regenerations with promising results [16]. Strattice was evaluated for seeding with rat MSCs [17]. Thus, the stem cell‐seeded biologic graft can be used as a tendon wrap or a skin regeneration material. A study was performed to compare the survival and proliferation of stem cells via bioluminescent imaging. The use of biologic graft

submucosa (SIS) treated with peracetic acid/ethanol (90%

crosslinked with hexamethylene diisocyanate (HMDI)

collagen, 5–10% lipids)

(97% collagen, 2% elastin)

collagen II)

Alloderm Human skin (decelluarized)

Restore 10‐layered porcine small intestine

BioBlanket Crosslinked porous bovine dermis

Strattice Porcine acelluar dermal matrix

**Table 2.** An example list of biologic ECM grafts with potential for tendon/skin regeneration.

TissueMend Noncrosslinked fetal biovine dermal matrix [19]

Permacol or Zimmer collagen repair patch Porcine acelluar dermis treated with trypsin, solvent, and

Cuffpatch EDC‐crosslinked matrix from eight layers of treated SIS

OrthoADAPT Crosslinked equine pericardium (90% collagen I, 10%

GraftJacket Human skin

human rotator cuff [14].

266 Biomaterials in Regenerative Medicine

(5 mm × 2 mm) collagen sponge scaffolds (Zimmer Dental) were used for the culture of (BM) MSCs. Cell‐seeded scaffolds were placed in culture dishes and incubated for 2 hours in a minimum volume of growth medium, after which more medium was applied to submerge the scaffolds. After an additional 24‐hour culture, cells seeded in scaffolds were treated with 10 ng/mL of recombinant (BMP) 12 for 12 hours. The medium was then replaced with fresh growth medium and scaffolds were either cultured for an additional 7 days or immediately implanted into partial calcaneal tendon defects in rats. It was shown that after 21 days, the BMP12‐treated, collagen‐cell scaffold results in robust formation of tendon‐like tissue [22]. Similarly, a collagen carrier combined with ASDCs showed they did not improve the biome‐ chanical properties of the tendon‐to‐bone healing. However, the ADSCs group showed less inflammation, which may lead to a more elastic repair and less scarred healing in a rat model.

#### *2.3.2. Integra bilayer*

Integra® bilayer wound Matrix (LifeSciences Corp., Plainsboro, New Jersey) is a dermal acellular analog composed of bovine collagen type I crosslinked with glycosaminoglycans. Importantly, inclusion of WJ‐MSC into Integra induced significant up‐regulation of prototypical angiogenic and healing factors, stimulating pleiotropic aspects of neovascularization in experimental set‐ tings of angiogenesis, without altering the inflammatory response in the animals, thus dem‐ onstrating their potential benefit in therapeutic care of wounds and skin grafts [23]. Recently, Integra dermal matrix scaffold engineered with adult mesenchymal stem cells and platelet‐rich plasma was investigated *in vitro* and demonstrated promising preliminary results [24].

#### **2.4. Synthetic biopolymer matrix**

Biopolymer matrix such as collagen, gelatin, hyaluronic acid, chitosan, silk, cellulose are the essential ECM components of human, animal, or plant cells. They have showed great promise for attachment, proliferation, and differentiation of stem cells and have been explored for tendon and skin regeneration.

#### *2.4.1. Collagen*

Collagen monomer solution can be extracted from skin of fetal bovine calf in close herd. Collagen‐based biomaterial has been widely used for tendon/skin regeneration. For tendon application, it is highly desirable to prepare aligned collagen fiber scaffolds. Anisotropic col‐ lagen biomaterial can be prepared by directional freeze drying [25], electrospinning, or a novel process called electrochemical process. Electrochemically aligned collagen fiber and skin substrate have been coupled with stem cells for both tendon [26] and skin regeneration [27], respectively. Collagen carrier/gel [28] or collagen combined with stem cell accelerated the wound healing in healing‐impaired db/db mice [29].

and proliferation of the mesenchymal stromal cells *in vitro*. *In vivo* assays performed with mesenchymal stromal cells grown on the surface of the dense membranes showed accelera‐

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A thiol‐ene Michael‐type addition was utilized for rapid encapsulation of MSCs within a gela‐ tin/PEG biomatrix according to Eq. (1). The MSCs/gelatin/PEG biomatrix was applied as a provisional dressing to full‐thickness wounds in Sprague‐Dawley rats. Biomatrix resulted in attenuated immune cell infiltration, lack of foreign giant cell (FBGC) formation, accelerated wound closure and re‐epithelialization, as well as enhanced neovascularization and granula‐

Gelatin ‐ PEG ‐ SH + PEG‐diacrylate → Gelatin ‐ PEG hydrogel (1)

Electrospun nanofiberous scaffolds prepared from silk fibroin protein were seeded with bone marrow‐derived mesenchymal stem cells (MSCs) and epidermal stem cells (ESCs). The con‐ structs were evaluated for wound re‐epithelization, collagen synthesis, as well as the skin appendages regeneration. It was shown that both the transplantation of MSCs and ESCs could significantly accelerate the skin re‐epithelization, stimulate the collagen synthesis. Furthermore, the regenerative features of MSCs and ESCs in activating the blood vessels and hair follicles for‐ mation, respectively, were suggested [35]. Combination of silk with collagen or poly(lactic‐co‐ glycolic acid) (PLGA) [18] and stem cells were evaluated in a rabbit tendon defect model [36].

A stroma skin substitute was first generated by using a mixture of human fibrin obtained from frozen human plasma and 0.1% agarose. An average of 250,000 cultured skin fibro‐ blasts were added to 5 mL of the mixture immediately before inducing the polymerization of the artificial stroma on Transwell (Corning Enterprises, Corning, NY, http://www.corning. com) porous inserts. Once the stromas jellified, human umbilical cord Wharton's jelly stem cells (HWJSCs) were seeded on top of the skin artificial stromas and cultured. It was demon‐ strated that this 3D bioactive scaffold can stratify and form epithelial cell‐like layers and well‐ formed cell‐cell junctions [37]. The authors also demonstrated similar strategy can be used for regeneration of oral mucosa using human Wharton's jelly stem cells (HWJSCs)‐seeded

Sodium CMC was evaluated as a substrate for ADSCs and implanted in adult male Wistar rats. CMC at 10 mg/mL associated with ADSCs increased the rate of cell proliferation of the granulation tissue and epithelium thickness when compared to untreated lesions (Sham).

tion of wound healing in Wistar rats [33].

tion tissue formation by 7 days [34].

*2.4.6. Gelatin/PEG*

*2.4.7. Silk*

*2.4.8. Fibrin‐agarose*

fibrin‐agarose‐mucosal fibroblasts.

*2.4.9. Sodium carboxymethylcellulose (CMC)*

#### *2.4.2. Pullulan‐collagen hydrogel*

A biomimetic pullulan‐collagen hydrogel was used to create a functional biomaterial‐based stem cell niche for the delivery of MSCs into wounds. Murine bone marrow‐derived MSCs were seeded into hydrogels and compared to MSCs grown in standard culture conditions. Hydrogels induced MSC secretion of angiogenic cytokines and expression of transcription factors associated with maintenance of pluripotency and self‐renewal. MSC‐seeded hydro‐ gels showed significantly accelerated healing and a return of skin appendages [30].

#### *2.4.3. Gelatin*

Gelatin is denatured collagen. Human ADSCs laden gelatin microcryogels (GMs) were evalu‐ ated *in vitro* as a stem cell carrier. The cell phenotype markers, stemness genes, differenti‐ ation, secretion of growth factors, cell apoptosis, and cell memory were compared against cells without a carrier. The priming effects of GMs on upregulation of stemness genes and improved secretion of growth factors of hASCs were demonstrated.

#### *2.4.4. Hyaluronic acid (HA)*

Hyaluronic acid (HA) is a nonsulfated, linear polysaccharide with the repeating disaccharide, β‐1,4‐D‐glucuronic acid ‐ β‐1,3‐N‐acetyl‐D‐glucosamine (Mw: 100–8000 KDa). It is an ECM component. ADSCs combined with hyaluronic acid (ADSC‐HA) dermal filler were implanted in rats and compared with HA alone. It was demonstrated that ADSC‐HA has better filling effects than HA alone. A total of 70% of stem cells remain in the injection site after 3 months. These suggested stem cells have the potential to improve the esthetic effects and longevity of dermal fillers [31].

#### *2.4.5. Chitin/Chitosan*

Chitosan‐hyaluron membrane: Hsu et al. investigated adult ADSCs spheroids combined with a chitosan‐hyaluron membrane and showed biomaterial combined with stem cells promoted wound healing in a rat skin repair model [32]. Dense and porous chitosan‐xanthan mem‐ branes seeded with multipotent mesenchymal stromal cells were evaluated for the treatment of skin wounds. The membranes showed to be nonmutagenic and allowed efficient adhesion and proliferation of the mesenchymal stromal cells *in vitro*. *In vivo* assays performed with mesenchymal stromal cells grown on the surface of the dense membranes showed accelera‐ tion of wound healing in Wistar rats [33].

#### *2.4.6. Gelatin/PEG*

*2.4.1. Collagen*

268 Biomaterials in Regenerative Medicine

*2.4.3. Gelatin*

*2.4.4. Hyaluronic acid (HA)*

dermal fillers [31].

*2.4.5. Chitin/Chitosan*

Collagen monomer solution can be extracted from skin of fetal bovine calf in close herd. Collagen‐based biomaterial has been widely used for tendon/skin regeneration. For tendon application, it is highly desirable to prepare aligned collagen fiber scaffolds. Anisotropic col‐ lagen biomaterial can be prepared by directional freeze drying [25], electrospinning, or a novel process called electrochemical process. Electrochemically aligned collagen fiber and skin substrate have been coupled with stem cells for both tendon [26] and skin regeneration [27], respectively. Collagen carrier/gel [28] or collagen combined with stem cell accelerated

A biomimetic pullulan‐collagen hydrogel was used to create a functional biomaterial‐based stem cell niche for the delivery of MSCs into wounds. Murine bone marrow‐derived MSCs were seeded into hydrogels and compared to MSCs grown in standard culture conditions. Hydrogels induced MSC secretion of angiogenic cytokines and expression of transcription factors associated with maintenance of pluripotency and self‐renewal. MSC‐seeded hydro‐

Gelatin is denatured collagen. Human ADSCs laden gelatin microcryogels (GMs) were evalu‐ ated *in vitro* as a stem cell carrier. The cell phenotype markers, stemness genes, differenti‐ ation, secretion of growth factors, cell apoptosis, and cell memory were compared against cells without a carrier. The priming effects of GMs on upregulation of stemness genes and

Hyaluronic acid (HA) is a nonsulfated, linear polysaccharide with the repeating disaccharide, β‐1,4‐D‐glucuronic acid ‐ β‐1,3‐N‐acetyl‐D‐glucosamine (Mw: 100–8000 KDa). It is an ECM component. ADSCs combined with hyaluronic acid (ADSC‐HA) dermal filler were implanted in rats and compared with HA alone. It was demonstrated that ADSC‐HA has better filling effects than HA alone. A total of 70% of stem cells remain in the injection site after 3 months. These suggested stem cells have the potential to improve the esthetic effects and longevity of

Chitosan‐hyaluron membrane: Hsu et al. investigated adult ADSCs spheroids combined with a chitosan‐hyaluron membrane and showed biomaterial combined with stem cells promoted wound healing in a rat skin repair model [32]. Dense and porous chitosan‐xanthan mem‐ branes seeded with multipotent mesenchymal stromal cells were evaluated for the treatment of skin wounds. The membranes showed to be nonmutagenic and allowed efficient adhesion

gels showed significantly accelerated healing and a return of skin appendages [30].

improved secretion of growth factors of hASCs were demonstrated.

the wound healing in healing‐impaired db/db mice [29].

*2.4.2. Pullulan‐collagen hydrogel*

A thiol‐ene Michael‐type addition was utilized for rapid encapsulation of MSCs within a gela‐ tin/PEG biomatrix according to Eq. (1). The MSCs/gelatin/PEG biomatrix was applied as a provisional dressing to full‐thickness wounds in Sprague‐Dawley rats. Biomatrix resulted in attenuated immune cell infiltration, lack of foreign giant cell (FBGC) formation, accelerated wound closure and re‐epithelialization, as well as enhanced neovascularization and granula‐ tion tissue formation by 7 days [34].

$$\text{Gelatin -PEG} \cdot \text{SH} + \text{PEC-diacrylate} \rightarrow \text{Gelatin -PEC hydrogen} \tag{1}$$

#### *2.4.7. Silk*

Electrospun nanofiberous scaffolds prepared from silk fibroin protein were seeded with bone marrow‐derived mesenchymal stem cells (MSCs) and epidermal stem cells (ESCs). The con‐ structs were evaluated for wound re‐epithelization, collagen synthesis, as well as the skin appendages regeneration. It was shown that both the transplantation of MSCs and ESCs could significantly accelerate the skin re‐epithelization, stimulate the collagen synthesis. Furthermore, the regenerative features of MSCs and ESCs in activating the blood vessels and hair follicles for‐ mation, respectively, were suggested [35]. Combination of silk with collagen or poly(lactic‐co‐ glycolic acid) (PLGA) [18] and stem cells were evaluated in a rabbit tendon defect model [36].

#### *2.4.8. Fibrin‐agarose*

A stroma skin substitute was first generated by using a mixture of human fibrin obtained from frozen human plasma and 0.1% agarose. An average of 250,000 cultured skin fibro‐ blasts were added to 5 mL of the mixture immediately before inducing the polymerization of the artificial stroma on Transwell (Corning Enterprises, Corning, NY, http://www.corning. com) porous inserts. Once the stromas jellified, human umbilical cord Wharton's jelly stem cells (HWJSCs) were seeded on top of the skin artificial stromas and cultured. It was demon‐ strated that this 3D bioactive scaffold can stratify and form epithelial cell‐like layers and well‐ formed cell‐cell junctions [37]. The authors also demonstrated similar strategy can be used for regeneration of oral mucosa using human Wharton's jelly stem cells (HWJSCs)‐seeded fibrin‐agarose‐mucosal fibroblasts.

#### *2.4.9. Sodium carboxymethylcellulose (CMC)*

Sodium CMC was evaluated as a substrate for ADSCs and implanted in adult male Wistar rats. CMC at 10 mg/mL associated with ADSCs increased the rate of cell proliferation of the granulation tissue and epithelium thickness when compared to untreated lesions (Sham). CMC is capable to allow the growth of ADSCs and is safe for this biological application up to the concentration of 20 mg/mL. These findings suggest that CMC is a promising biomaterial to be used in cell therapy [38].

#### **2.5. Synthetic nondegradable polymer‐based biomaterial**

**Table 3** summarizes common nondegradable polymer used as a biomaterial or as a modifier of ECM‐based biopolymer for skin and tendon tissue engineering applications. The incorporation of such polymer allows biopolymer (collagen, gelatin, HA) to be able to gelled at physiological condition (e.g., under UV/blue light, ambient temperature, or room temperature aqueous free radical initiator. Stem cells can be directly encapsulated inside the substrate during gel formation.

angiogenesis. Moreover, USCs could secrete vascular endothelial growth factor(VEGF) and transforming growth factor (TGF)‐β1, and USC‐conditioned medium enhanced the migra‐ tion, proliferation, and tube formation of endothelial cells. This data suggested that USCs in combination with PCL/GT significantly prompted the healing of full‐thickness skin wounds in rabbits. Similarly, electrospun poly (L‐lactide‐co‐ε‐caprolactone)/poloxamer (PLCL/P123) scaffolds combined with ADSCs enhanced skin wound healing in a rat model [47]. Chitosan‐ crosslinked poly (3‐hydroxybutyrate‐co‐3‐hydroxyvalerate) (PHBV) was used to load unre‐ stricted somatic stem cells. The cell‐laden scaffold showed better results during the healing

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**PLGA PCL PHBV**

**Table 4.** Common synthetic, degradable polymer used for tendon/skin regeneration applications.

Biomaterials play an important role for attachment, survival, and function of stem cells. Many biomaterials are either used alone or as one component of the product for the regeneration of tendon/ligament/skin. Despite abundance of biomaterial developed, the optimal biomaterials that meet the structural, mechanical, functional requirement of tendon/ligament tissues to be regenerated remain a challenge. Novel biomaterial fabrication process, biomaterial design, and biomaterial synthesis toward tendon/ligament and skin regeneration are urgently needed.

[1] Uysal CA, et al. Adipose‐derived stem cells enhance primary tendon repair: Biomechanical and immunohistochemical evaluation. Journal of Plastic Reconstructive & Aesthetic

process of skin defects in rat models [48, 49].

Address all correspondence to<EMAIL_ADDRESS>
Surgery. 2012:**65**(12):1712‐1719

Southwest Research Institute, San Antonio, TX, USA

**3. Conclusion**

**Author details**

Xingguo Cheng

**References**

#### *2.5.1. Poly(NIPAM)‐based*

A biodegradable, multifunctional crosslinker and an n‐isopropylacrylamide (NIPAM)‐based, thermosensitive hydrogel was synthesized to carry BMSCs to treat diabetic skin ulcers. The crosslinker contains an arginylglycylaspartic acid(GRD)‐like motif that promotes cell attach‐ ment and differentiation of BMSCs. After hydrogel association with BMSCs treated the diabetic skin wound in mice, significantly greater wound contraction was observed in the hydrogel + BMSCs group. Histology and immunohistochemistry results confirmed that this treatment contributed to the rapid healing of diabetic skin wounds by promoting granulation tissue formation, angiogenesis, extracellular matrix (ECM) secretion, wound contraction, and re‐epithelialization. These results show that a hydrogel laden with BMSCs may be a promis‐ ing therapeutic strategy for the management of diabetic ulcers [39].

**Table 3.** Common synthetic, nondegradable polymer used as hydrogel or gel components.

#### **2.6. Polyester‐based degradable polymer biomaterial**

The structure of common polyester is shown in **Table 4**. Stem cell‐coated polyester suture was evaluated for tendon applications in a rat model [40]. Electrospun PLGA fiber may be more suit‐ able for tendon regeneration than film [41]. (PGA/PLA) fiber combined with ADSCs improve tendon in a rabbit tendon model [42]. Open cell PLGA seeded with stem cells produced more collagen type I [43]. Knited PLGA encapsulated with stem cell/alginate gel [44]. PGA sheet with MSCs were able to regenerate tendon‐bone insertions and the tendon in rabbit [45]. Electrospun polycaprolactone/gelatin (PCL/GT) membrane and human urine‐derived stem cells (USCs) were evaluated for skin wound healing in a rabbit model [46]. USCs‐PCL/GT‐treated wounds closed much faster, with increased re‐epithelialization, collagen formation, and

**Table 4.** Common synthetic, degradable polymer used for tendon/skin regeneration applications.

angiogenesis. Moreover, USCs could secrete vascular endothelial growth factor(VEGF) and transforming growth factor (TGF)‐β1, and USC‐conditioned medium enhanced the migra‐ tion, proliferation, and tube formation of endothelial cells. This data suggested that USCs in combination with PCL/GT significantly prompted the healing of full‐thickness skin wounds in rabbits. Similarly, electrospun poly (L‐lactide‐co‐ε‐caprolactone)/poloxamer (PLCL/P123) scaffolds combined with ADSCs enhanced skin wound healing in a rat model [47]. Chitosan‐ crosslinked poly (3‐hydroxybutyrate‐co‐3‐hydroxyvalerate) (PHBV) was used to load unre‐ stricted somatic stem cells. The cell‐laden scaffold showed better results during the healing process of skin defects in rat models [48, 49].

## **3. Conclusion**

CMC is capable to allow the growth of ADSCs and is safe for this biological application up to the concentration of 20 mg/mL. These findings suggest that CMC is a promising biomaterial

**Table 3** summarizes common nondegradable polymer used as a biomaterial or as a modifier of ECM‐based biopolymer for skin and tendon tissue engineering applications. The incorporation of such polymer allows biopolymer (collagen, gelatin, HA) to be able to gelled at physiological condition (e.g., under UV/blue light, ambient temperature, or room temperature aqueous free radical initiator. Stem cells can be directly encapsulated inside the substrate during gel formation.

A biodegradable, multifunctional crosslinker and an n‐isopropylacrylamide (NIPAM)‐based, thermosensitive hydrogel was synthesized to carry BMSCs to treat diabetic skin ulcers. The crosslinker contains an arginylglycylaspartic acid(GRD)‐like motif that promotes cell attach‐ ment and differentiation of BMSCs. After hydrogel association with BMSCs treated the diabetic skin wound in mice, significantly greater wound contraction was observed in the hydrogel + BMSCs group. Histology and immunohistochemistry results confirmed that this treatment contributed to the rapid healing of diabetic skin wounds by promoting granulation tissue formation, angiogenesis, extracellular matrix (ECM) secretion, wound contraction, and re‐epithelialization. These results show that a hydrogel laden with BMSCs may be a promis‐

The structure of common polyester is shown in **Table 4**. Stem cell‐coated polyester suture was evaluated for tendon applications in a rat model [40]. Electrospun PLGA fiber may be more suit‐ able for tendon regeneration than film [41]. (PGA/PLA) fiber combined with ADSCs improve tendon in a rabbit tendon model [42]. Open cell PLGA seeded with stem cells produced more collagen type I [43]. Knited PLGA encapsulated with stem cell/alginate gel [44]. PGA sheet with MSCs were able to regenerate tendon‐bone insertions and the tendon in rabbit [45]. Electrospun polycaprolactone/gelatin (PCL/GT) membrane and human urine‐derived stem cells (USCs) were evaluated for skin wound healing in a rabbit model [46]. USCs‐PCL/GT‐treated wounds closed much faster, with increased re‐epithelialization, collagen formation, and

to be used in cell therapy [38].

270 Biomaterials in Regenerative Medicine

*2.5.1. Poly(NIPAM)‐based*

**2.5. Synthetic nondegradable polymer‐based biomaterial**

ing therapeutic strategy for the management of diabetic ulcers [39].

**Table 3.** Common synthetic, nondegradable polymer used as hydrogel or gel components.

**Poly(NIPAM) PEGDA Poly(methyl acrylate) (PMA)**

**2.6. Polyester‐based degradable polymer biomaterial**

Biomaterials play an important role for attachment, survival, and function of stem cells. Many biomaterials are either used alone or as one component of the product for the regeneration of tendon/ligament/skin. Despite abundance of biomaterial developed, the optimal biomaterials that meet the structural, mechanical, functional requirement of tendon/ligament tissues to be regenerated remain a challenge. Novel biomaterial fabrication process, biomaterial design, and biomaterial synthesis toward tendon/ligament and skin regeneration are urgently needed.

## **Author details**

Xingguo Cheng Address all correspondence to<EMAIL_ADDRESS>Southwest Research Institute, San Antonio, TX, USA

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**Chapter 12**

**Provisional chapter**

**Hydrogels in Regenerative Medicine**

**Hydrogels in Regenerative Medicine**

DOI: 10.5772/intechopen.70409

Polymer scaffolds have many various applications in the field of tissue engineering, drug delivery, and implantation. They are applied as dispensing devices for bioactive molecules and as three-dimensional (3D) structures that provide stimulants that organize cells and direct desired original tissue formation. Hydrogels are preferred scaffolding material because they are structurally similar to the extracellular matrix of many tissues, often processed under mild conditions, and can be delivered in a minimally invasive manner. Hydrogel materials formed a group of polymeric materials. The hydrophilic structure allows them to hold large amounts of water in their three-dimensional backbone. As a result, hydrogels are used as scaffolding material for drug and growth factor transmission, tissue engineering modifications, and many other applications. In this chapter, we describe the physical and chemical structure of hydrogels, side groups, cross-linkings, swelling properties, types of polymers and fabrication methods, and application fields.

**Keywords:** hydrogel, polymers, drug delivery, tissue engineering

preferred as they promote cell attachment and proliferation.

hydrogels in regenerative medicine is aimed.

Yasemin Budama-Kilinc, Rabia Cakir-Koc, Bahar Aslan, Burcu Özkan, Hande Mutlu and

Yasemin Budama-Kilinc, Rabia Cakir-Koc, Bahar Aslan, Burcu Özkan, Hande Mutlu and

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

Hydrogels are polymer-based substances that contain high levels of water and show the physical characteristic of the extracellular matrix. In regenerative medicine, in terms of the material requirements for tissue scaffold or therapeutic transfer systems, the cell matrix is preferred due to its structural and compound-based similarities. Moreover, hydrogels are

In this chapter, compiling the studies about the preparation of hydrogels and application of

http://dx.doi.org/10.5772/intechopen.70409

Eslin Üstün

**Abstract**

**1. Introduction**

Eslin Üstün

**Provisional chapter**

## **Hydrogels in Regenerative Medicine**

**Hydrogels in Regenerative Medicine**

Yasemin Budama-Kilinc, Rabia Cakir-Koc, Bahar Aslan, Burcu Özkan, Hande Mutlu and Eslin Üstün Bahar Aslan, Burcu Özkan, Hande Mutlu and Eslin Üstün Additional information is available at the end of the chapter

Yasemin Budama-Kilinc, Rabia Cakir-Koc,

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.70409

#### **Abstract**

Polymer scaffolds have many various applications in the field of tissue engineering, drug delivery, and implantation. They are applied as dispensing devices for bioactive molecules and as three-dimensional (3D) structures that provide stimulants that organize cells and direct desired original tissue formation. Hydrogels are preferred scaffolding material because they are structurally similar to the extracellular matrix of many tissues, often processed under mild conditions, and can be delivered in a minimally invasive manner. Hydrogel materials formed a group of polymeric materials. The hydrophilic structure allows them to hold large amounts of water in their three-dimensional backbone. As a result, hydrogels are used as scaffolding material for drug and growth factor transmission, tissue engineering modifications, and many other applications. In this chapter, we describe the physical and chemical structure of hydrogels, side groups, cross-linkings, swelling properties, types of polymers and fabrication methods, and application fields.

DOI: 10.5772/intechopen.70409

**Keywords:** hydrogel, polymers, drug delivery, tissue engineering

## **1. Introduction**

Hydrogels are polymer-based substances that contain high levels of water and show the physical characteristic of the extracellular matrix. In regenerative medicine, in terms of the material requirements for tissue scaffold or therapeutic transfer systems, the cell matrix is preferred due to its structural and compound-based similarities. Moreover, hydrogels are preferred as they promote cell attachment and proliferation.

In this chapter, compiling the studies about the preparation of hydrogels and application of hydrogels in regenerative medicine is aimed.

## **2. Physical and chemical properties of hydrogels**

#### **2.1. Crosslinking and general characteristic of hydrogels**

Hydrogels are a three-dimensional (3D) polymeric network structure with the ability to absorb water and other biological fluids. The water content of hydrogels determines its specific physicochemical character: mildness, density, and low surface tension of hydrogels [1–4].

hydrophilicity of the polymer and thus increases the water absorption capacity. As a result of the electrostatic repulsive forces of these loads, the solubility and swelling ratios in the web structure are increased. In ionic hydrogels containing acid groups, swelling increases in basic medium, while hydrogels containing basic groups increase swelling in acidic media [9, 10].

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 279

Hydrogels are polymeric webs that absorb and retain water in large quantities. In the polymeric cannula, there are hydrophilic groups or areas that provide a hydrated hydrogel structure in an aqueous environment. As is the case with the term "network," cross-links must be

Physical gel called hydrogels is the result of cross-linking of ionic interaction and physical interactions such as hydrogen bonds, coordination bonds, and hydrophobic interactions [11]. These hydrogels alter the temperature, pH, or solvent composition to form a homogeneous solution and re-gel when they return to their initial conditions. Because of these behaviors, physical gel is also known as reversible gel [12]. They are generally unstable and mechanically weak [13].

Generally, ionic interactions are used to produce physically cross-linked hydrogels. For example, alginate hydrogels are formed with calcium ions. The physical cross-links are found between ionic interactions [14], crystallization [14], amphiphilic block and graft copolymers

The hydrogels in this group are called as irreversible because they are cross-linked by strong chemical bonds between their chains and are not redissolved by changing temperature, pH, or solvent composition. Chemical cross-linked nets have permanent junctions [18, 19]. Chemical cross-linking involves permanent covalent bonds that result in better mechanical strength and more stable hydrogels than physical cross-linking. Incorporation of chemical cross-linkers, however, can cause toxicity problems. Cross-linking of hydrogel can occur with some other method such as radical polymerization, chemical reaction, and high-energy irradiation.

Chemically cross-linked gels can be obtained by radical polymerization of polymers of low molecular weight drugs in the presence of cross-linking agents; poly(2-hydroxyethyl methacrylate) (pHEMA) is a well-known and frequently studied hydrogel system. This hydrogel was first described by Wichterle and Lim [19] and was obtained by polymerization of HEMA in the presence of a suitable cross-linking agent (e.g., ethylene glycol dimethacrylate). A wide

The covalent bonds between the polymer chains can be formed by reaction with complementary reactivity of the functional groups, formation of an amine carboxylic acid, or OH/

 reaction such as isocyanate or Schiff base formation. Chemical cross-linking bonds can occur with aldehydes, with additional reactions and condensation reactions. Water-soluble

variety of hydrogel systems have been synthesized using similar procedures [20].

present to prevent the hydrophilic polymer chains from dissolving in aqueous phase.

**2.3. Cross-linking methods in hydrogel fabrication**

[15], hydrogen bonds [16], and protein interactions [17].

*2.3.1. Physically cross-linked hydrogel*

*2.3.2. Chemical cross-linked hydrogels*

NH3

Hydrogels can be classified according to their preparation methods, ionic charges, physical structures, and cross-linking conditions. According to physical structures, hydrogels can be divided into some groups that are called amorphous hydrogels, semi-crystalline hydrogels, hydrogen-bonded hydrogels, physical hydrogels, and chemical hydrogels.

Various cross-linking approaches, including chemical and physical, have been used to create polymer networks and to protect 3D structures in aqueous media. Physical interactions between the polymer chains in the physically cross-linked gels prevent hydrogel breakdown, while the covalent bonds between the polymer chains in the chemically cross-linked gel form a constant hydrogel.

Physically cross-linked hydrogels are formed by changes in environmental conditions (pH, temperature, and ionic interactions), hydrogen bonds, and protein interactions. Chemically cross-linked gels are obtained by radical polymerization, chemical reactions, energy irradiation, and enzymatic cross-linking [5].

Hydrogels can also have many physical forms: solid mold forms (such as soft contact lenses), pressed powder matrices (pills or capsules taken orally), microparticles (bio-structured carriers or wound treatments), coatings (in implants or in catheters, in pills or capsules, or in capillary electrophoresis), membranes or disks (such as two-dimensional electrophoresis gels), encapsulated solute (in osmotic pumps), and liquids (gel forms that are heated or cooled) [6].

#### **2.2. Hydrogels side groups**

Hydrogels are also called "hydrophilic polymers" because of their water-like properties. Hydrogels have functional groups, such as **─**SO<sup>3</sup> H, **─**COOH, **─**CONH<sup>2</sup> , **─**OH, and **─**NH<sup>2</sup> , which provide hydrophilic character to the polymer chains in their networks. Some of these functional groups (**─**SO<sup>3</sup> and **─**COO hydrogel nets) are negatively charged in aqueous media and can adsorb metal ions through electrostatic interactions [7].

Depending on their chemical and physical structure, there are three types of hydrogels: neutral hydrogels, ionic hydrogels, and interconnected network structures. Neutral (nonionic) hydrogels are homopolymeric or copolymeric hydrogels which do not have charged groups in their structure. Ionic hydrogels, also known as polyelectrolytes, are prepared from ionically charged monomers. These hydrogels are termed cationic and anionic hydrogels, as their monomeric charge is positive and negative [8]. The presence of charged groups in the main chain of ionic hydrogels enhances their susceptibility to stimuli. The presence of ionizable functional groups such as carboxyl and sulfonic acid or ammonium salt increases the hydrophilicity of the polymer and thus increases the water absorption capacity. As a result of the electrostatic repulsive forces of these loads, the solubility and swelling ratios in the web structure are increased. In ionic hydrogels containing acid groups, swelling increases in basic medium, while hydrogels containing basic groups increase swelling in acidic media [9, 10].

#### **2.3. Cross-linking methods in hydrogel fabrication**

Hydrogels are polymeric webs that absorb and retain water in large quantities. In the polymeric cannula, there are hydrophilic groups or areas that provide a hydrated hydrogel structure in an aqueous environment. As is the case with the term "network," cross-links must be present to prevent the hydrophilic polymer chains from dissolving in aqueous phase.

#### *2.3.1. Physically cross-linked hydrogel*

**2. Physical and chemical properties of hydrogels**

hydrogen-bonded hydrogels, physical hydrogels, and chemical hydrogels.

Hydrogels are a three-dimensional (3D) polymeric network structure with the ability to absorb water and other biological fluids. The water content of hydrogels determines its specific physicochemical character: mildness, density, and low surface tension of hydrogels [1–4]. Hydrogels can be classified according to their preparation methods, ionic charges, physical structures, and cross-linking conditions. According to physical structures, hydrogels can be divided into some groups that are called amorphous hydrogels, semi-crystalline hydrogels,

Various cross-linking approaches, including chemical and physical, have been used to create polymer networks and to protect 3D structures in aqueous media. Physical interactions between the polymer chains in the physically cross-linked gels prevent hydrogel breakdown, while the covalent bonds between the polymer chains in the chemically cross-linked gel form

Physically cross-linked hydrogels are formed by changes in environmental conditions (pH, temperature, and ionic interactions), hydrogen bonds, and protein interactions. Chemically cross-linked gels are obtained by radical polymerization, chemical reactions, energy irradia-

Hydrogels can also have many physical forms: solid mold forms (such as soft contact lenses), pressed powder matrices (pills or capsules taken orally), microparticles (bio-structured carriers or wound treatments), coatings (in implants or in catheters, in pills or capsules, or in capillary electrophoresis), membranes or disks (such as two-dimensional electrophoresis gels), encapsulated solute (in osmotic pumps), and liquids (gel forms that are heated or cooled) [6].

Hydrogels are also called "hydrophilic polymers" because of their water-like properties.

which provide hydrophilic character to the polymer chains in their networks. Some of these

Depending on their chemical and physical structure, there are three types of hydrogels: neutral hydrogels, ionic hydrogels, and interconnected network structures. Neutral (nonionic) hydrogels are homopolymeric or copolymeric hydrogels which do not have charged groups in their structure. Ionic hydrogels, also known as polyelectrolytes, are prepared from ionically charged monomers. These hydrogels are termed cationic and anionic hydrogels, as their monomeric charge is positive and negative [8]. The presence of charged groups in the main chain of ionic hydrogels enhances their susceptibility to stimuli. The presence of ionizable functional groups such as carboxyl and sulfonic acid or ammonium salt increases the

H, **─**COOH, **─**CONH<sup>2</sup>

and **─**COO hydrogel nets) are negatively charged in aqueous media

, **─**OH, and **─**NH<sup>2</sup>

,

**2.1. Crosslinking and general characteristic of hydrogels**

a constant hydrogel.

278 Biomaterials in Regenerative Medicine

tion, and enzymatic cross-linking [5].

Hydrogels have functional groups, such as **─**SO<sup>3</sup>

and can adsorb metal ions through electrostatic interactions [7].

**2.2. Hydrogels side groups**

functional groups (**─**SO<sup>3</sup>

Physical gel called hydrogels is the result of cross-linking of ionic interaction and physical interactions such as hydrogen bonds, coordination bonds, and hydrophobic interactions [11]. These hydrogels alter the temperature, pH, or solvent composition to form a homogeneous solution and re-gel when they return to their initial conditions. Because of these behaviors, physical gel is also known as reversible gel [12]. They are generally unstable and mechanically weak [13].

Generally, ionic interactions are used to produce physically cross-linked hydrogels. For example, alginate hydrogels are formed with calcium ions. The physical cross-links are found between ionic interactions [14], crystallization [14], amphiphilic block and graft copolymers [15], hydrogen bonds [16], and protein interactions [17].

#### *2.3.2. Chemical cross-linked hydrogels*

The hydrogels in this group are called as irreversible because they are cross-linked by strong chemical bonds between their chains and are not redissolved by changing temperature, pH, or solvent composition. Chemical cross-linked nets have permanent junctions [18, 19]. Chemical cross-linking involves permanent covalent bonds that result in better mechanical strength and more stable hydrogels than physical cross-linking. Incorporation of chemical cross-linkers, however, can cause toxicity problems. Cross-linking of hydrogel can occur with some other method such as radical polymerization, chemical reaction, and high-energy irradiation.

Chemically cross-linked gels can be obtained by radical polymerization of polymers of low molecular weight drugs in the presence of cross-linking agents; poly(2-hydroxyethyl methacrylate) (pHEMA) is a well-known and frequently studied hydrogel system. This hydrogel was first described by Wichterle and Lim [19] and was obtained by polymerization of HEMA in the presence of a suitable cross-linking agent (e.g., ethylene glycol dimethacrylate). A wide variety of hydrogel systems have been synthesized using similar procedures [20].

The covalent bonds between the polymer chains can be formed by reaction with complementary reactivity of the functional groups, formation of an amine carboxylic acid, or OH/ NH3 reaction such as isocyanate or Schiff base formation. Chemical cross-linking bonds can occur with aldehydes, with additional reactions and condensation reactions. Water-soluble polymers owe their solubility properties to the presence of functional groups (mainly **─**OH, COOH, NH) which can be used for hydrogel formation.

*2.5.1. Natural and synthetic polymers*

tides, and leucine [37] zipper;

lose, and gelatin/alginate [39]; and

among these three main types.

*2.5.2. Biodegradable synthetic polymers*

gels by chemical or physical cross-linking [51, 52].

• DNA [40].

Natural polymers have been used as scaffolds for tissue engineering as natural hydrogels, regarding their biological compatibility, natural biodegradability, and critical biological func-

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 281

• proteins such as collagen, silk, gelatin, fibrin, lysozyme, Matrigel™, and genetically engineered proteins [34–36] like calmodulin (a calcium-binding protein), elastin-like polypep-

• protein/polysaccharide hybrid polymers like collagen/HA, laminin/chitosan, fibrin/cellu-

Compared to natural polymers, synthetic polymers have more reproducible physical and chemical properties that are critical for the construction of scaffolding. At present, synthetic polymers have emerged as an important alternative option in the manufacture of hydrogel tissue-engineered skeletons, since they can be molecularly modified by block structures, molecular weights, mechanical strength, and biodegradability [41–43]. Synthetic polymers used in the preparation of synthetic hydrogels can be classified into three main types: nonbiodegradable [44, 45], biodegradable [46], and bioactive polymers [47]. In the following section, the biodegradable synthetic polymers are disclosed because of its widespread use from

Biodegradability is one of the most important aspects of scaffolds for tissue engineering. It is highly desirable to ensure that the biodegradation rate coincides with new tissue regeneration at the defect site [48, 49]. Many polymers, proteins, cellulose, starch, and chitin created in the environment are limited to making hydrogel skeletons that are biodegradable but have their own biodegradability and mechanical properties. Most common biodegradable synthetic polymer groups contain poly (lactic acid) (PLA), poly(ε-caprolactone) (PCL), poly(glycolic acid) (PGA) and copolymers [50]. They can be used to modify hydrophilic polymers such as PEG to form acrylate macromers or amphiphilic polymers to produce biodegradable hydro-

The mechanical stability of the gel is an important feature when a scaffold is designed. The strength of the hydrogels can be increased by incorporating cross-linking agents, comonomers, because the higher cross-linking degree provides more fragility and less elasticity.

Biodegradable synthetic hydrogels include various vinylic monomers such as 2-hydroxyethyl methacrylate (HEMA), N-isopropylacrylamide (NIPAM), 2-hydroxypropyl methacrylate (HPMA), acrylamide (AAm), acrylic acid (AAc) or macromers [53–55] N,N′-methylenebis and methoxyl poly(ethylene glycol) (PEG), monoacrylate (mPEGMA or PEGMA), and diacrylate (PEGDA) with cross-linkers such as ethylene glycol diacrylate (EGDA), and these

• polysaccharides like hyaluronic acid (HA), dextran, agarose, and chitosan [38].

tions. There are four main types of natural polymers, including:

High-energy radiation, especially gamma and electron beams, can be used to polymerize unsaturated compounds. This means that water-soluble polymers derivatized with vinyl groups can be converted into hydrogels by high-energy irradiation [21]. Hydrogels may also be obtained by radiation-induced polymerization of a mixture consisting of a monofunctional acrylate (e.g., acryloyl-L-proline methyl ester) and a suitable cross-linker [22]. Moreover, highenergy irradiation cross-links water-soluble polymers without additional groups. During irradiation of aqueous solutions of polymers (gamma or electron transfer), the radicals are generally present in the polymer chain, e.g., homolytic cleavage of C–H bonds.

#### **2.4. Water content and resources**

Hydrogels can generally hold 20–100 times water and their own weight depends on their structures [23]. If hydrogels can hold higher than 100%, they are called superabsorbent hydrogels. According to the water contents, hydrogels can be categorized as low swelling grade hydrogels (20–50%), medium swelling grade hydrogels (50–90%), high swelling grade hydrogels (90–99.5%), and superabsorbent hydrogels (>99.5%) [24, 25].

High water content hydrogels are usually more useful in the medical field because of permeability and biocompatibility [26]. Synthesis of biodegradable hydrogels with both good mechanical properties and high water content is still a major challenge, because as the amount of water increases, the mechanical properties of hydrogel become weak [27].