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Keywords:

  • biocompatibility;
  • biodegradable;
  • catalysis;
  • electron beam curing;
  • monolith;
  • reaction injection molding;
  • ROMP

Abstract

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. POLYMERIC MATERIALS FOR TISSUE ENGINEERING
  5. SUMMARY AND OUTLOOK
  6. Acknowledgements
  7. REFERENCES AND NOTES
  8. Biographical Information

The state of the art in polymeric materials for tissue engineering as well as the needs and concerns for future medical applications are outlined and discussed and brought into relation to recent developments in polymer chemistry. Particularly, the recent developments in micro- and nano-structured polymeric monoliths designed for these purposes will be discussed. © 2009 Wiley Periodicals, Inc. J Polym Sci Part A: Polym Chem 47: 2219–2227, 2009


INTRODUCTION

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. POLYMERIC MATERIALS FOR TISSUE ENGINEERING
  5. SUMMARY AND OUTLOOK
  6. Acknowledgements
  7. REFERENCES AND NOTES
  8. Biographical Information

Osseous Tissue Engineering

The lack of sources for autologous bone transplantation has lead to the development of different strategies for bone replacement, one presented by biological and alloplastic bone substitute materials with and without the additional application of growth factors (bone regeneration). Until now, however, the autologous bone has been considered to be the absolute and unrivalled gold standard. Alternatively, to accomplish complete osseous defect regeneration, tissue engineering,1–8 i.e., the implantation of scaffolds seeded in vitro with autologous cells, has been proposed. The use of bone substitute materials has the intention to provide a framework for the ingrowing of mesenchymal cells with osteoblastic differentiation9 in terms of osteoconductivity or osteoinductivity.10 Thereby it is aimed that the primarily avital scaffold material will be replaced by functional, vital bone tissue. Of particular importance within that context are polymer-composite materials,11 which are either based on biodegradable or non-degradable polymers. These provide both the (cell-loaded) polymeric scaffold as well as the inorganic material, for example, hydroxyapatite (HA) that facilitates the formation of osseous tissue. In first clinical applications in preprosthetic augmentation procedures of the maxilla,12 however, ossification was often incomplete, although the sinus lift defect actually offers ideal conditions for bone transplantation and regeneration.13 In fact, so far the results have not been better than those obtained by the use of alloplastic bone substitute materials alone, such as HA. Possible reasons for the failure are manifold and include insufficiencies of the scaffold materials in terms of chemistry and mechanics, the lack of osteogenic stimuli and the problem of nutrition and vascularization. Collagenous products lack mechanical stability, besides the unsolved problem of senzitation. In addition, inorganic materials with high mechanical stability such as HA,14, 15 tricalciumphosphate (TCP),16, 17 and other ceramics, are replaced only insufficiently by vital bone and are often integrated by fibrous, scar-like tissue. This is not surprising, since the exceptional properties of bone material are the consequence of very subtle (sub-) structures that strongly depend on mineral morphologies and orientation as well as on crystallinity,18 which are all to be controlled during growth, i.e., in vivo19 and not in vitro.

Soft Tissue Engineering

Apart from osseous tissue, there is a strong need for soft tissue augmentation in various surgical fields for adipose engineering. Beside its essential metabolic functions, adipose tissue provides the shape and volume of the outer contour of the body and preserves the mobility of tissue layers. Consequently, soft tissue augmentation has a strong impact in many instances of plastic and reconstructive surgery and especially facial surgery. Soft tissue augmentation is thus necessary in case of congenital disorders, after tumor ablation as well as in esthetic or reconstructive surgery.

The problems with adipose tissue transplantation are well known: Because of its high metabolic activity, adipose tissue cannot be transplanted freely in larger amounts and needs vascular supply.20 Transplantation of larger amounts of adipose tissue results in colliquation necrosis and development of oilagenous cysts. This circumstance has lead to the development of lipofilling techniques,21 which provide persistent tissue augmentation by the injection of very small adipose tissue fragments. Although it is thereby possible to increase the cell survival rate of transplanted fat and to increase the persistent volume in the grafted area,22 long-term results are still unpredictable. Further disadvantages are the limits of the volume attained and the necessity for multiple-stage injection.

Transplantation of undifferentiated cells with or without adipogenic commitment might overcome these problems. It is known, that progenitor and precursor cells have much lower metabolic requirements than mature adipocytes. Implantation of precursor cells might result in persistent tissue augmentation by differentiation of the transplanted cells to mature adipocytes.

For tissue augmentation to be used for esthetic improvement, so-called “dermal fillers” are often used to treat contour defects. However, common problems associated with resorbable fillers are sensitation, allergic reactions, and the necessity of repeated treatments. Permanent fillers on the other hand can lead to inflammation and scarification, are difficult to remove and are therefore not approved by the FDA. Current approaches for adipose tissue engineering use gels as a carrier material. Porous alginates,23 synthetically derived non-resorbable products like poly(methylmethacrylate) (PMMA) and HA as well as hyaluronan-based (HYADD3) gels and composite materials are used.24 The latter were most favorable used in pigs, however, a considerable shrinkage of the implanted volume was observed within 4 weeks.25 Consequently, stability of the engineered tissue volume remains a problem. Even worse, collagenous microparticles have additional disadvantages: Their resorption is too fast (even during the in vitro phase) resulting in a loss of cells. Any modification (e.g., by crosslinking procedures) is not in conformity with GMP requirements. A further aspect is the provision of nutrition and vasculature, which is essential for adipose tissue engineering.

From a medical point of view, autologous lipofiling appears to be the best choice to augment soft tissues of the face as compared with the injection of allogenic or alloplastic materials.26 Although it is thereby possible to increase the persistent volume in the grafted area, long-term results are unpredictable. Only small volumina can be transplanted successfully, further disadvantages are the limits of the attained volume and the necessity for multiple-stage injection. Within this context, adipose tissue engineering might be able to enlarge the potential and the versatility of autologous soft tissue augmentation. The advantages over existing methods would be: (a) persistent augmentation after single or few injections, (b) no side effects due to sensitizing agents, (c) possibility for later corrections, (d) cost-effectiveness as compared with current dermal filers.

Alternatively, soft tissue augmentation can be necessary also in oncologic patients for esthetic and also for functional reasons. After surgical tumor treatment in the head and neck region, soft tissue augmentation is often needed and remains a challenge especially after irradiation.

POLYMERIC MATERIALS FOR TISSUE ENGINEERING

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. POLYMERIC MATERIALS FOR TISSUE ENGINEERING
  5. SUMMARY AND OUTLOOK
  6. Acknowledgements
  7. REFERENCES AND NOTES
  8. Biographical Information

Requirements

Apart from the requirement of being affordable and FDA approvable, some additional hard criteria need to be met by a material to be feasible for applications in the areas of cell cultivation, differentiation, or tissue engineering.27 First, it needs to possess a highly porous structure. These pores can form during the degradation process or can already be implemented into the polymeric structure. In fact, the need to provide porous materials has soon been realized. Recently, Ratner underlined the necessity of specifically engineered porosities that do not only allow for but in fact can induce vascularization and less fibrotic healing.28 In orthopedic applications, porous metal frameworks29 are used that allow for the ingrowing of osseous tissue, resulting in an optimized bonding of the implant to the bone host. Differences in porosity can be used to tune the elastic moduli, allowing for a better match between the old and new material. In addition, the material and all its pores needs to be hydrophilic to promote diffusion of water-based body liquids. For the final vascularization, pore diameters around 150 μm appear favorable. Only pores of such diameters allow for both the substantial ingrowing of the cells and for vascularization at a comparably late stage of cell proliferation. Second, the matrix itself needs to be biocompatible. On the one hand, this requires a surface chemistry that favors cell contact and adhesion. On the other hand, the polymer may not contain or, in course of the degradation process, form any compounds that exhibit any cell toxicity. In addition, the matrix should display reasonable mechanic strength, and finally, be biodegradable. The fragments that form in course of this degradation process should have molecular weights <40,000 g/mol in order be released from the body via the kidneys. However, a major advantage of monolithic supports is related to the fact that they may be created in virtually any form via a simple molding process. This allows for creating entire parts of the body. Finally, the scaffold needs to maintain its shape. Within that context, it is evident that the requirements concerning elastic and storage modulus, compressibility, etc., in soft tissue engineering applications are different from those in osseous tissue applications.

Polymers

In terms of polymers, particularly those based on poly(hydroxyalkanoate)s, for example, poly(glycolic acid) (PGA), poly(L-lactic acid) (PLLA), poly(3-hydroxybutyrate) (PHB), poly(3-hydroxybutyrate-co-3-hydroxyvalerate), poly(β-propiolactone) (PPL), poly(ε-caprolactone) (PCL), poly(ethylene succinate) (PES), poly(butylene succinate) (PBS), are biodegradable and thus often used in tissue engineering both in form of the polymer itself as well as in form of polymer/inorganic phase composite materials (vide supra).30 Apart from these compounds, however, polymeric scaffolds and composite materials based on poly(ester urethane)urea, lysine diisocyanate-glucose, poly(propylene fumarate), poly(glycerol sebacate), poly(carbonate), poly(phospazene), trimethylene carbonate, and others need to be mentioned.31

Processing Techniques

It would certainly by far exceed the scope of this article to profoundly discuss all the basics, advantages, disadvantages, promises, and draw-backs of the different techniques available. Therefore, the most important fabrication techniques for the creation of porous, 3D materials shall only be mentioned briefly. Thus, the use of porous membranes,32 sol-gel-derived glasses, metals and metal alloys;27 foamed33 or electrospinning-derived polymers has been reported. Porosity may also be generated by polymer templating or rapid prototyping techniques such as 3D plotting and 3D printing,34–42 stereolithography, selective laser sintering, and fused deposition modeling.27, 43–49 The latter allows for the use of numerous polymers including biodegradable ones such as PCL, PLLA, as well as of composite materials, for example, PP-TCP, PCL-HA, PCL-TCP, PLLA-TCP, or chitosan-HA. More recently, alternative approaches, for example, the assembling of photolithographically-derived cell-laden hydrogels50 have been reported. Nevertheless, most of these techniques represent a bottom-up approach, where porosity is generated during an assembling process.

Porous Monolithic Materials

In search of alternative scaffolds suitable for tissue engineering, we focused on monolithic supports. Monolithic polymeric materials are characterized by a unitary porous structure with interconnected large pores, usually in the low micrometer range, and may be synthesized within the confines of the compartment in which they are to be used at a later stage.51, 52 This means that porosity is generated during polymer synthesis and can in fact be controlled.53–56

To obtain monoliths within these definitions, a polymerization setup has to be created in which phase separation of an increasingly crosslinked matrix from the polymerization medium, usually consisting of a good and a poor polymer solvent serving as the micro- and macro- porogen, is faster than gelation. According to the Flory-Huggins theory,57–59 large, positive interaction parameters between the polymer and the solvent or solvent mixture are required to induce phase separation. It means that at least one non-solvent for the polymer that forms has to be used. In addition, molding processes are feasible, significantly reducing restrictions in shape. During the last 15 years, polymer-based monolithic supports have been introduced into separation science as well as into organo- and biocatalysis.60, 61 Major contributions have been made by the group of Fréchet and Svec, who initially developed radical polymerization-based systems for these purposes.51, 62, 63 Our group contributed to this area by developing a ring-opening metathesis polymerization- (ROMP-) based approach, which allowed for a one-pot synthesis of functionalized monolithic systems. Using ROMP-based synthetic protocols, a large variety of monolithic systems have been realized.52, 54, 64, 65 Recently, we elaborated an additional synthetic pathway based on the electron-beam (EB) triggered free radical polymerization-based synthesis of highly porous monolithic materials.66–70 These two approaches and the adherent basics and concepts shall be outlined in more detail.

EB-Triggered, Free Radical Polymerization-Derived Monoliths

In the presence of a suitable initiator, the free radical polymerization of acrylates, methacrylates, acrylamides, and other vinylic compounds can be triggered thermally or photochemically. This approach has been widely used for the synthesis of monolithic supports.52 As an alternative, the initiator-free, free radical polymerization of (meth-) acrylates is a well-known procedure, particularly in printing and coating technology. The thickness of substrate that can be converted into a polymer (means “cured”) strongly depends on the energy of the EB source. Thus, a 10 MeV source penetrates roughly 5 mm of glass or steel and up to 7 cm of a polymerization solution. This is by far enough for the synthesis of 3D monolithic structures for tissue engineering, since most scaffolds do not exceed 7 cm in diameter. Another advantage of the EB-based technique is that it requires no heating. This and the fact that EB-guns can be run in a pulsed mode allows for extending the time of polymerization, thus effectively dissipating the heat of polymerization. Thus ultimately results in homogeneous monolithic structures. Finally, it should also be mentioned that the EB based process entails the simultaneous sterilization of the matrix by the electron beam.

For our purposes, 2-methylidene-4-phenyl-1,3-dioxolane (10 wt %) was used as monomer, trimethylolpropane triacrylate (21 wt %) was chosen as a crosslinker. A mixture of toluene (microporogen), 2-propanol and dodecan-1-ol (macroporogens, 10:30:30 wt %) was chosen as micro and macroporogen. The resulting crosslinked copolymer displayed a porous structure with pores up to 50 μm. Its formation and chemical structure is shown in Scheme 1 (left). Since the crosslinker is an ester and the dioxolane is transformed into an ester in course of the free radical polymerization process,71 the entire polymer is based on a polyester structure that may well be expected to be biodegradable. A representative scanning electron microscopic picture is shown in Figure 1.

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Figure 1. Porous, EB-initiated free radical polymerization-derived (left) and ROMP-derived monolithic scaffolds. The scale bar represents 100 μm (left) and 400 μm (right). The scale bar in the right insert corresponds to 4 μm.

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Scheme 1. Structure of polyester-based, EB-initiated free radical polymerization-derived monoliths (left); 7-oxanorborn-2-ene-based, ROMP-derived monoliths (right).

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For biocompatibility studies, tests of the in vitro growth and for cell differentiation studies, human adipose tissue derived stem cells (ATSCs)72 were used. ATSCs have advantages from a clinical point of view because of their abundant availability as compared to bone marrow-derived stem cells and are likewise suited for osseous and adipose tissue engineering. Cultivation of ATSCs on sterilized polyester-based monolithic supports revealed a rapid cell growth on the surface of the materials, suggesting biocompatibility. However, almost no ingrowth of the cells into the monolithic structure was observed. Even worse, after 20 days, the cells disintegrated from the support, indicating insufficient adhesion. These findings are believed to be related to a too hydrophobic character of the polymer itself and a surface micro-structure of the support, which additionally disfavors cell contact over a long period of time. All this is in fact supported by the very high water contact angle (142°) that was observed with these materials and that must be a result of both the chemistry and surface structure of the polymer.73–75 As a consequence of this unsatisfying situation, current work focuses on the use of more polar monomers and crosslinkers.

Ring-Opening Metathesis Polymerization-Derived Monoliths

As a result of the impressive accomplishments in catalyst development and mechanistic understanding, ring-opening-metathesis polymerization (ROMP) nowadays holds a strong position in polymer chemistry.76–80 Particularly, Grubbs-type initiators, for which an appealing synthesis had been developed and which became commercially available short time after,79–83 are characterized by a high robustness towards protic functionalities. This permits the polymerization of functionalized, protic norborn-2-enes and 7-oxanorborn-2-enes even in aqueous media. Most importantly, synthesis is facilitated, making the use of glove box or Schlenk conditions unnecessary. Finally, the development of Grubbs-type initiators based on the Ru(IMesH2)(2-Br-pyridine)2(alkylidene)-motive did not only increase their reactivity but also made them comparable to Schrock-type initiators in terms of control over polymerization and livingness (IMesH2 = 1,3-dimesitylimidazolin-2-ylidene).84 In view of these developments, we investigated the general applicability of ROMP to the synthesis of monoliths and found it in fact suitable for our purposes. Thus, we established both NBE54, 55, 64, 85–87 and cis-cyclooctene-based,56, 88 ROMP-derived monoliths for numerous applications. Because ROMP is a living polymerization technique and polymerization as performed at low temperature (0 °C), there is excellent control over the polymerization and phase separation process resulting in an unprecedented reproducibility of synthesis and control over porosity.88

In an effort to create polymeric scaffolds that promote cell adhesion and allows for the ingrowing of cells, we turned to a ROMP-based protocol based on highly hydrophilic monomers. Thus, by using a 20:20 wt % mixture of norborn-2-ene (NBE) and pentaglycerol bis(7-oxanorborn-5-ene-2-ylcarboxylate) acrylate (PGBA) in a microporogen (toluene) and a macroporogen (2-propanol, 5:10 wt % ratio), monolithic structures were realized with the aid of RuCl2Py2(IMesH2)(CHPh) (0.03 wt %, Py = pyridine) and low amounts of additional Py (0.071 wt %) as regulator.89 As has been shown for related systems,56 the presence of pyridine in amounts as low as 70 ppm allows for the additional tuning of the monoliths' structure by affecting the polymerization kinetics, which is first-order in Py for RuCl2Py2(IMesH2)(CHPh). It should be mentioned that the acrylate group in PGBA acts as a chain transfer agent, thus significantly reducing the average chain lengths of the norborn-2-ene derived polymer blocks. Scheme 1 (right) illustrates the process of monolith synthesis; a typical structure with pores around 200 μm is shown in Figure 1. These large pores translate into an interparticle-derived porosity of 77%, as determined by inverse-size exclusion chromatography.90 In view of any intoxicating effects of catalyst residues, a careful washing of the monolithic structure with a mixture of dimethylsulfoxide, ethyl vinyl ether, and THF was carried out and allowed for the virtually quantitative removal of the transition metal-based initiator, resulting in final ruthenium concentrations <0.1 ppm (i.e., below the limit of detection of inductively coupled plasma optical emission spectroscopy, ICP-OES). Upon storage in a buffer system, the initial water contact angle of 142° rapidly decreased within 13 weeks. In fact, no reliable values could be determined due to fast moistening, i.e., spreading. This is attributed to both the highly hydrophilic character of PGBA and the high propensity towards oxidation of the poly(NBE) blocks, which experience oxidation in allylic position.

In a next step, the ROMP-derived scaffolds were cultivated with ATSCs. Within 12 days, the number of cells quadruplated at least at the surface. No additional compounds such as fibronectine, RGD-peptides or matrix factors, factors (aside from differentiation factors or supplements) commonly used to facilitate or trigger cell growth, were necessary. To study cell ingrowth and differentiation on the scaffolds in a three-dimensional setting, cultivation was performed under dynamic conditions in rotating culture containers. Important enough, cells grew into the pores and partially showed even full ingrowth and penetration of the scaffold. Finally, after growth, these cells were then successfully subjected to both osteogenic and adipocytic differentiation. Differentiation of ATSCs into either adipocytes or osteoblasts was initiated by addition of isobutylmethylxanthine or ß-glycerolphosphate and dexamethasone, respectively. Figure 2(A) shows adipocytes as identified by the formation of large univacuolar cells. Figure 2(B) shows mineralizations by the cells grown within the monolithic material after 6 weeks. As can be seen, biomineralization was effective with these materials.

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Figure 2. Histomicrographs of monoliths seeded with adipose tissue derived stem cells after adipogenic (left) and osteogenic differentiation (right). (A) Low power magnification of scaffold cross section with cell ingrowth between the scaffold material after adipogenic differentiation. Fat cells with nuclei (“F”) are also visible in the center of the cross section of the scaffold (HE). (B) Von Kossa stain shows massive mineralizations (arrows, black-brown) after osteogenic differentiation within the monolithic scaffold (counterstaining of cell nuclei with HE). The scaffold material is invaded by cellular ingrowth.

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Mechanical Properties of Monoliths

The typical Young's modulus of ROMP-derived monolithic materials is around 1 MPa, the compressive strength is 7.4 MPa, and the Martens Hardness is 0.3 N/mm2. These values are a direct consequence of the highly porous structure and the low rigidity of the monomers used. Despite their low mechanical strength, they were found suitable for adipose tissue engineering. For osseous tissue engineering, however, a significant increase, for example, a factor of 1000 in the Young's modulus appears most favorable.

SUMMARY AND OUTLOOK

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. POLYMERIC MATERIALS FOR TISSUE ENGINEERING
  5. SUMMARY AND OUTLOOK
  6. Acknowledgements
  7. REFERENCES AND NOTES
  8. Biographical Information

In summary, we have added monoliths as another tool to the armor of fabrication methods of scaffolds for tissue engineering. The data obtained so far are promising, however, not more. Many things have to be accomplished, such as increasing the mechanical strength, tuning biodegradation, optimizing cell contact in EB triggered, free radical polymerization derived monoliths and many more. Thus, carrier materials need not only to be biocompatible but should also enable both the growth and the proliferation of mesenchymal stem cells, i.e., guide adipogenic differentiation by a specific nano- and micro-patterned surface,91–94 preferable in the absence of any additional chemical stimulus. This appears accomplishable, since apart from other factors, cell differentiation, proliferation is strongly governed by mechanical stress or, more generally, interaction of cells with its surrounding matrix.95–102 In addition, scaffolds must possess a mechanical strength sufficient for the corresponding application and must cause only minimal in vivo reactions (if any) and provide a controlled degradation process. The final goal is to be able to implant viable micro-constructs, which are able to differentiate the corresponding progenitor cells towards mature tissue and lead to a persistent tissue augmentation not only in healthy patients, but also in irradiated sites. If necessary, the polymerization mixtures used for monolith synthesis should also contain survival and angiogenic factors, which must in turn also be addressed by both catalyst and synthesis issues. Apart from polymer chemistry and surface science, methods for the augmentation of soft tissues based on adipose tissue-derived progenitor cells with injectable carrier materials have to be developed in a clinically relevant setting, for example, for patients who have experienced ablative tumor surgery and radiation therapy.

Acknowledgements

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. POLYMERIC MATERIALS FOR TISSUE ENGINEERING
  5. SUMMARY AND OUTLOOK
  6. Acknowledgements
  7. REFERENCES AND NOTES
  8. Biographical Information

Support provided by B. Frerich, Universitäts-Klinik für Mund-, Kiefer- und Plastische Gesichtschirurgie, Universität Leipzig, Germany, is gratefully acknowledged.

REFERENCES AND NOTES

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. POLYMERIC MATERIALS FOR TISSUE ENGINEERING
  5. SUMMARY AND OUTLOOK
  6. Acknowledgements
  7. REFERENCES AND NOTES
  8. Biographical Information

Biographical Information

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. POLYMERIC MATERIALS FOR TISSUE ENGINEERING
  5. SUMMARY AND OUTLOOK
  6. Acknowledgements
  7. REFERENCES AND NOTES
  8. Biographical Information
Thumbnail image of

Michael R. Buchmeiser was born in Linz, Austria. He received his PhD degree in Organometallic Chemistry in 1993 from the University of Innsbruck, Austria. He was awarded an “Erwin Schrödinger Fellowship” and spent 1 year at the MIT within the group of Prof. Richard R. Schrock (Chemistry Nobel Prize 2005). In 1998, he finished his “Habilitation” in Macromolecular Chemistry (University of Innsbruck) where he then held a Faculty Position as Associate Professor from 1998–2004. He received the “1998 Professor Ernst Brandl Research Award,” the “START Award-2001” as well as the “Novartis Award 2001.” He was offered Faculty Positions (Full Professor of Polymer Chemistry) at the University of Halle (Germany) in 2004, at the University of Leoben (Austria) in 2005, and at the TU Dresden in 2007, which he all declined. Instead, since December 2004, he holds a Faculty Position (C-4 Professor) at the University of Leipzig, Germany. In addition, since February 2005, he is Vice Director and Member of Board at the Leibniz Institute of Surface Modification (IOM), Leipzig, Germany. Starting December 2009, he will take over a Faculty Position (Full Professor) at the University of Stuttgart (Germany) and be the Director of the Institute of Textile Chemistry and Chemical Fibers. So far, he has published more than 200 research papers and has filed approximately 20 patents.