Kenneth E. Gonsalves, Craig R. Halberstadt, Cato T. Laurencin, Lakshmi S. Nair. Biomedical Nanostructures

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CHAPTER 15

Nanostructures for Tissue

Engineering/Regenerative Medicine

SYAM P. NUKAVARAPU, SANGAMESH G. KUMBAR,

LAKSHMI S. NAIR, and CATO T. LAURENCIN

15.1 INTRODUCTION

15.1.1 Tissue Engineering/Regenerative Medicine

Tissue and organ failure has been a major health problem in the United Stastes

and worldwide. Autografting and allografting are the two main approaches

currently used to repair or replace damaged or lost tissue and organs. However,

limited availability and donor site morbidity are the problems associated with

autografting, while disease transmission and foreign body response are

associated with allografting. In addition, there is also a major concern about

the growing number of patients waiting for an organ replacement. According

to a recent report, in the United States alone 89,000 patients are waiting for an

organ transplant and 17 people die every day waiting for a transplant of a vital

organ such as a heart, liver, kidney, pancreas, lung, or bone marrow [1]. The

viable solution to this problem is tissue engineering/regenerative medicine,

which is deﬁned as the application of biological, chemical, and engineering

principles toward the repair, restoration, or regeneration of living tissues using

biomaterials, cells, and factors, alone or in combination [2].

Researchers so far have identiﬁed three major strategies to regenerate tissues

using tissue engineering principles [3]. The ﬁrst approach involves injecting

tissue-speciﬁc cells at the injury site. This is a noninvasive process and practiced

in the repair of minor injuries, however not useful for regenerating complex

tissues involving more than one cell type with a speciﬁc architecture. A second

approach is guided tissue engineering, where biodegradable scaffolds made of

BiomedicalNanostructures, EditedbyKennethE.Gonsalves,CraigR.Halberstadt,CatoT.Laurencin,

and Lakshmi S. Nair

377

poly(L-lactic acid) (PLLA), polyglycolic acid (PGA), poly(D,L-lactide-co-

glycolide) (PLGA), polycaprolactone (PCL), polyphosphazenes, and so on

are used for guiding cellular in growth. The material selection is based on the

desired application and is designed to permi t scaffold disintegration by the time

cells invade and completely repair the injured tissue.

Third and the most important strategy utilizes biodegradable scaffolds in

combination with growth factors and tissue-speciﬁc cells or stem cells that

ultimately differentiate into the required cell type. This approach is further

divided into two kinds, one is a closed system and the other is an open system.

In a closed system, cell-seeded scaffolds are encapsulated using a thin

semipermeable membrane that acts as a barrier an d protects cells from the

host immune response while allowing nutrient transport and waste removal

essential for cell survival. This approach is quite similar to extracorporeal

device technology where particular organ functionality is temporarily restored.

The open system is the most popular approach, where scaffolds are precultured

in vitro before implantation. Based on the requirement, growth factors that are

helpful in tissue regeneration are either released from the scaffold or delivered

through programming cells via gene therapy [4]. Biodegradable scaffolds in

combination with cells and growth factors have been used successfully to

regenerate various tissues like skin, bone, cartilage, blood vessels, and heart

valves. This method is further improved by replacing static culture methods

with dynamic culture systems using bioreactors. These advanced culture

systems not only mitigate nutrient transport limitations but also provide

mechanical stimula tion to seeded cells in the form of ﬂuid shear stress [5].

Although tissue engineering approach has seen early success in the regenera-

tion of sim ple tissues like skin and cartilage, there are however biological and

engineering concerns to be addressed before the approach becomes a successful

clinical alternative for regenerating complex tissues. In order to become a

practical and successful procedure, scaffold selection, dosage of cells, and the

amount and the kind of growth factors have to be optimized based on the

particular application and the relative health status of the patient.

15.1.2 Scaffolds for Tissue Engineering

Tissue engineering involves assembling relevant cells into the required three-

dimensional (3D) organs or tissue. Cells lack this ability to form a 3D tissue in

culture and end up forming 2D layers with no anatomical features. The three-

dimensionality can be achieved by seeding cells onto an artiﬁcial structure known

as a scaffold that is capable of supporting 3D tissue formation, which is an

important component in tissue engineering [6]. The required features of a

scaffold are biocompatibility and bioactivity, biodegradability or bioresorb-

ability, mechanical compatibility, interconnected porosity, and the nontoxic

nature of the degradation products. Biocompatibility is deﬁned as supporting the

appropriate cellular activity without eliciting any undesirable effects in those cells

and the host [7, 8]. Bioactivity refers to having appropriate surface chemistry that

378 BIOMEDICAL NANOSTRUCTURES

favors cell attachment, proliferation, and differentiation [9, 10]. Controlled

biodegradability is crucial because the rate of degradation of the scaffold should

be in line with the neotissue formation, since scaffolds need to be absorbed by the

surrounding tissues without the necessity of a surgical removal. High pore

volume and an adequate pore size are necessary for cell inﬁltration and also for

effective nutrient transport and waste removal. For example, scaffolds with 30%

pore volume and 150 mm median pore size performed well for bone tissue

regeneration [2]. Matching mechanical properties of the intended tissue (high

modulus for bone and low modulus materials for vascular tissue engineering) is

also critical for successful tissue engineering.

Synthetic polymers such as PGA, PLA, PLGA, PCL, polyphosphazenes, and

polyanhydrides, and naturally derived proteins and carbohydrate polymers such

as chitosan, alginate, and hyluronic acid are the most commonly used

biodegradable polymers for tissue engineering scaffold applications [11, 12].

Over the years, several techniques have been developed to fabricate synthetic and

natural polymeric materials into 3D porous scaffolds. Some of the conventional

scaffold fabrication techniques include solvent casting and particulate leaching

[13], gas foaming [14], freeze drying [15], 3D printing (solid freeform fabrication)

[16], and microsphere sintering [17]. It has been demonstrated that most of the

3D porous scaffolds were capable of supporting cellular adhesion and

proliferation, but lack the ability to simulate the extracellular matrix (CM)

like environment that is always in close proximity with cells in all the tissues.

Therefore, the best alternative is nanofeatured scaffolds that not only mimic the

ECM conditions but also show higher reactivity for proteins and ultimately

enhances cell adhesion and thus regenerative capacity. Techniques like

electrospinning [18], temperature-induced phase separation [19], molecular self-

assembly [20], and surface patterning [21] are currently in practice for generating

scaffolds with nano- and micron-size features for better mimicking the native

ECM. In this chapter, we focus on the importance of nanofeatured scaffolds and

their fabrication using various techniques. Further emphasis is also made on in

vitro and in vivo studies where the tissue regenerative capacity is enhanced with

these novel scaffolds.

15.1.3 Nanofeatures of ECM

ECM is a self-assembled nanoﬁbrillar network compo sed of complex

biomacromolecules that surround and support cells in tissues [22]. ECM is

present in interstitial complex (IC) and basement membrane (BM) along with

cells as shown in Fig. 15.1a. ECM is constructed with different classes of

biomolecules such as structural proteins (collagen and elastin) and specialized

proteins (ﬁbrillin, ﬁbronectin, and laminin) that help to structurally connect,

bind, integrate cells to form tissues, and also provide a surface for cellular

adhesion and migration. Proteoglycans are glycoaminoglycans like chondroitin

sulfate and heparin sulfate linked with serine like core proteins that provide

chemical cues and regulate cell growth, differentiation, and metabolic activity.

NANOSTRUCTURES FOR TISSUE ENGINEERING/REGENERATIVE MEDICINE 379

The interstitial complex is mostly present in connective tissues like bone,

cartilage, and ligaments , and present in association with a wide variety of cells

and the above-mentioned proteins. BM supports a single layer of endothelial,

epithelial, and some mesenchymal cells in many organs (Fig. 15.1b). BM’s

function is as an active barrier for cell inﬁltration and migration, and also acts

as a reservoir for a variety of growth factors.

ECM and BM are comprised of ﬁbrils made of proteins such as collagen,

elastin, and ﬁbronectin as shown in Fig. 15.1c. These ﬁbrils are of nanosize and

much smal ler than micrometer-sized cells. The major component in the ECM is

collagen, whi ch is the ﬁbrous backbone of the ECM an d is at least available in

28 forms in our body. The abundant collagen is mostly in a ﬁbril structure

composed of 300-nm-long, 1.5-nm- diameter rodlike protei ns consisting of

three coiled chains. Each chain consists of 1050 amino acids that are wound

around in a characteristic right-handed triplex. Further, interaction of triple

helices of collagens results in the formation of ﬁbrils that are roughly 50 nm in

FIGURE 15.1 (a) SEM image of a chick embryo cornea showing epithelial cells, basal

lamina, and network of collagen ﬁbrils. Molecular structure of (b) basal lamina and

Trelstad, (b) John Wiley & Sons Inc., and from (c) Garland Science Publishing).

380

BIOMEDICAL NANOSTRUCTURES

diameter. The ﬁbril packing is such that adjacent triple helix is displaced

approximately 67 nm to yield a banded structure. Another important

component in the ECM is ﬁbronectin that acts as a glue between cells and

the ECM by attaching cells to a variety of ECMs. Fibronectins (FNs) are

dimers of two similar peptides that are 6070 nm long and 23 nm thick. FN

contains nanodomains with an arginine glycineasparticserine sequence

that selectively exhibits high afﬁnity for a particular substrate such as collagen,

ﬁbrin, and cell surface receptors. In addition, other proteins that are present in

the ECM are also of nanosize and inﬂuence cell behavior at various levels.

15.2 NANOFIBROUS SCAFFOLDS

Cells within a tissue and organ are organized and encapsulated by the

nanoﬁbrous ECM. For example, a SEM of rat cornea tissue is shown in

Fig. 15.2a that is composed of ﬁbroblasts surrounded by ECM nanoﬁbrils. A

similar cellular arrangement is observed when cells are cultured in vitro on

nanoﬁbrous scaffolds (Fig. 15.2b). The tissue-like cellular organization observed

with nanoﬁbrous scaffolds prompted researchers to look for more sophisticated

methods to fabricate nanoﬁbrous structures that can better mimic the ECM and

serve as potential scaffolds for tissue regeneration. Electrospinning, phase

separation, and molecular self-assembly techniques that are currently popular

for fabricating nanoﬁber-based scaffolds are described below.

15.2.1 Electrospinning

Electrospinning is a facile, efﬁcient, and inexpensive technique for fabricating

ultraﬁne ﬁbers. A typical setup for electrospinning is shown in Fig. 15.3a.

During electrospinning, a high electric potential in the order of few kilovolts is

FIGURE 15.2 SEM of (a) the rat cornea tissue where ﬁbroblasts are surrounded

by extracellular matrix ﬁbrils (reproduced from Reference [22] with permission, from

Garland Science, publishing) and (b) human dermal ﬁbroblasts cultured in vitro on PCL

nanoﬁbers (reproduced from Appl Biochem Biotechnol 2005;125:147158 with

permission, Humana Press Inc.)

NANOSTRUCTURES FOR TISSUE ENGINEERING/REGENERATIVE MEDICINE 381

applied on the pendent drop of the polymer solution or melt held by surface

tension. The applied ﬁeld causes charge separation and hence causes repulsion

within the polymer drop sitting at the nozzle of the spinneret. When the

opposing electrostatic force overcomes the surface tension of the polymer

solution, a thin polymer jet ejection initiates from the spinneret. The emerging

jet undergoes a series of bending and stretching instabilities and becomes

elongated until it is deposited on the grounded collector. Solvent evaporation

occurs during the process, and eventually a polymer jet is deposited into long

and ultrathin ﬁbers with diameters ranging from tens of nanometers to a few

micrometers. Recently, this approach has draw n much interest within the tissue

engineering community. A number of synthetic and natural polymers have

been spun into a variety of structures having ﬁbers randomly oriented,

oriented, coreshell structured, porous, and drug loaded that are suitable for

tissue regeneration and drug delivery applications (Fig. 15.3a) [23–27].

Electrospun nonwoven nanoﬁber matrices are promi sing as scaffolds for

tissue engineering since the ﬁber diameters are in line with most of the natural

ECMs, and the resulting high surface area might allow a high percentage of

FIGURE 15.3 Nanoﬁbrous scaffolding methodologies; (a) Typical setup for electro-

spinning and the resulting nonoriented and oriented nanoﬁber scaffolds (reproduced

from Reference [33] with permission, from Elsevier). (b) Steps involved in the

temperature-induced phase separation process, and the obtained nanoﬁbrous and

foams with spherical pore shape (reproduced from Reference [19] with permission, from

Elsevier). (c) Self-assembly process and the derived individual nanoﬁbers and the

consolidated nanoﬁbrous scaffold (reproduced from Reference [49] with permission

from AAAS).

382

BIOMEDICAL NANOSTRUCTURES

cellular attachment. In addition to the nanoﬁbrous structure, the matrices also

exhibit porosities greater than 90 % with adequate pore size for cell migration

[27]. Thus, nanoﬁber matrices provide more structural space for cell

accommodation and make the nutrient and metabolic was te exchange process

efﬁcient between the scaffold and the surroundings.

Electrospinning technique for creating tissue engineering scaffolds has been

demonstrated and patented by Laurencin and Ko [28]. These authors

demonstrated for the ﬁrst time the formation of electrospun PLGA ultraﬁne

ﬁbers with morphological similarity to the ECM of natural tissue with a

diameter ranging from 500 to 800 nm with porosity greater than 90% [27]. In

vitro studies with mouse ﬁbroblasts further indicated good cellular adherence

and spreading onto PLGA nanoﬁbers. Based on these results, it was

hypothesized that 25100 mm pore diameter that was observed is adequate

for human mesenchymal stem cell migration and tissue ingrowth.

Incorporation of drugs into these nanoﬁbers has also been explored by these

authors for wound healing applications [29]. In a recent study, Nair et al. from

the same group developed polyphosphazene-based nanoﬁber scaffolds for

tissue engineering applications. These novel scaffolds support the adhesion and

proliferation of coronary artery endothelial cells and osteoblast-like MC3T3-

E1 cells [30]. Currently, we are focusing on developing various polypho-

sphazene ﬁber scaffolds alone or in combination with PLGA for bone tissue

engineering and wound dressing applications.

Many other resear ch groups have also successfully electrospun biodegrad-

able polyesters such as PLA, PGA, and their copolymer PLGA into nonwoven

ﬁber scaffolds [31, 32]. PLLA is electrospun into aligned and random nanoﬁber

mats with varied ﬁber diameters ranging from 300 to 700 nm. Ramakrishna

et al. studied the cellular compatibility of PLLA nanoﬁbers using neural stem

cells [33]. It was noticed that cells adhered and proliferated on aligned and

randomly oriented PLLA nanoﬁbrous scaffolds . Other synthetic biodegradable

polymers like PCL and PCL blends with polyorthoesters were also studied for

their spinnability and cell interaction. Fetal bovine chondrocytes seeded on

PCL nanoﬁber scaffolds, with an average ﬁber size of 700 nm, were able to

maintain the chondrocytic phenotype during 3 weeks of culture period [34].

Geng et al. demonstrated better cell attachment and proliferation with human

umbilical vein endothelial cells onto 50 : 50 poly(

L-lactic acid-co-e-caprolac-

tone) (PLCL) ﬁbers of 300 nm diameter when compared to the matrices with

7-mm ﬁbers [35]. The observed better cellular performance on these nanoﬁber

matrices is presumably due to the nanoﬁber structure that mimics the ECM.

Various natural polymers have also been considered for electrospinning

because of their prov en biocompatibility and biofunctionality. Natural

polymers such as collagen, ﬁbrinogen, alginate, hyaluronic acid, chitosan,

and starch were electrospun alone or in combination with synthetic polymers

to improve scaffold performance [36, 37]. Nanoﬁber scaffolds of ECM

polymers may have added advantage of resembling not only the size scale but

also the chemical and biological functions of the ECM. Collagen I and III were

NANOSTRUCTURES FOR TISSUE ENGINEERING/REGENERATIVE MEDICINE 383

electrospun using 1,1,1,3,3,3-hexaﬂuro-2-propanol (HFT) as a solvent that

created type I collagen nanoﬁbers with diameters in the range of 100 nm and

type III collagen in the range of 250 nm [38]. The fabricated nanoﬁber scaffolds

exhibited good cytocompatibility and proliferation with smooth muscle cells,

and also showed cell inﬁltration into the collagen network. Efforts were made

to improve the mechanical properties of the collagen nanoﬁ ber scaffolds by

PCL coating. Coating imparted the right mechanical strength while retaining

the collagen regions on the surface that maintained cell attachment and

proliferation of human dermal ﬁbroblasts for the purpose of dermal tissue

engineering [39]. Even though electrospinning is a proven mild process,

biological integrity of some polymers has been found to be lost during

electrospinning. Such an adverse affect is presumably due to the harsh organic

solvents like HFT used for electrospinning that could potentially denature the

protein structures. Therefore, electrospinning of natural polymers is limited

and still not versatile as compared to synthetic polymers.

15.2.2 Phase Separation

Polymer phase separation has been known for many years and is widely used

for developing membranes for nonmedical applications [40]. Recently,

thermally induced phase separation (TIPS) has become a popular technique

for fabricating porous and nanofeatured 3D scaffolds for tissue engineering

applications [41]. In this method, a polymer is ﬁrst dissolved in a suitable

solvent at high temperature. Lowering the solution temperature induces

liquidliquid or solidliquid phase separation into two phases, one with a

high polymer concentration and the other with a low polymer co ncentration.

Subsequent solvent removal by exchange, evaporation, or sublimation from

low polymer phase results in pores, whereas the high polymer phase solidiﬁes

into 3D interconnected ﬁbrous networks as seen in Fig. 15.3b. Scaffold

parameters such as porosity and ﬁber size can be controlled by varying the

solvent, the polymer concentration, and the temperature for any polymer

system.

Polymer phase separation has gained immense interest in recent days for

developing tissue engineering scaffolds for two reasons. The ﬁrst reason is that

controlling scaffold porosity is ideal for supporting cell proliferation and 3D

tissue formation. Second is the ability to form nanoﬁbrous 3D structure,

which mimics the ﬁbrous 3D structure of natural collagen. Figure 15.3b shows

SEM of highly porous scaffolds from aliphatic polyesters such as PLLA,

poly(

D-L-lactic acid) (PDLLA), and PLGA produced by phase separation

[42–44]. However, the na noﬁbrillar structure is only observed with the highly

crystalline polymer PLLA . Ma et al. extensively studied the effect of gelation

temperature, polymer concentration, and solvent s used on porosity and ﬁber

diameters of scaffolds produced using PLLA. Combining the polymer with

porogens further modiﬁed this process and achieved macroporous networks in

scaffolds with nanoﬁbrous pore walls [41].

384 BIOMEDICAL NANOSTRUCTURES

15.2.3 Molecular Self-Assembly

Molecular self-assembly is the spontaneou s organization of molecules into

structurally well-deﬁned and stable arrangements through weak and non-

covalent interactions. Such interactions include hydrogen bonding, ionic

bonding, hydrophobic interactions, van der Waals interactions, and water-

mediated hydrogen bonding. Recently, molecular self-assembly has become

one of the powerful methods to produce nanoﬁbrous scaffolds for tissue

engineering with properties similar in scale and chemistry to that of the natural

ECM.

Proteins, peptides, and lipid molecules are commonly used as building

blocks for forming nanofeatured scaffolds for tissue engineering applications.

Nature has already used these molecules to produce various struc tures like

collagen and keratin through molecular self-assembly [45]. For the ﬁrst time,

Zhang et al. produced 3D peptide scaffolds utilizing naturally occurring amino

acids by exposing the self-assembling peptide ((Ala-Glu-Ala-Glu-Ala-Lys-Ala-

Lys)

(EAK16)) to physiological media or a salt solution [46]. Peptides with

alternating hydrophilic and hydrophobic amino acid moieties aggregated

depending upon salt concentration and pH to form stable scaffolds consisting

of interwov en nanoﬁbers. This method has opened a new area of research

where the peptide solution forms into a scaffold in vivo with ECM

functionalities that would help in repair and regeneration of damaged or lost

tissues. Based on this concept BD biosciences has commercialized a product

‘‘PuraMatrix peptide hydrogel’’ with ﬁber diameters of 10 nm and pores of

50200 nm. This matrix has shown in vitro differentiation of hepatocyte

progenitor cells into hepatocytes [47], neurons from hippocampal slices [48],

and neurite out growth from PC12 cells [96].

Recently, Stupp et al. designed peptide-amphiphile (PA) nanocylinders

adopting a composite approach with nonpolar hydrocarbon tails and

hydrophilic peptide heads as seen from Fig. 15.3c [49]. The nanocylinder

formation is induced through a pH change from 8 to 4. The produced ﬁbers

(diameter of 7.6 1 nm) were cross-linked by oxidation to achieve structural

integrity. Phosphorylated serine incorporation within the peptide end leads to

the calciophilic nature of the scaffold with the ability to form hydroxyapatite

mineral coating that promise a new approach for bone tissue engineering [50].

In vitro studies with different cell types on 3D peptide scaffolds exhibited

interesting functional behavior such as good proliferation, functional

differentiation, and active migration [51]. In a separate study with bovine

chondrocytes encapsulated in a peptide scaffold, Kisiday et al. showed the

extensive production of their own ECM (different types of collagen with

glycosaminoglycans) while maintaini ng their phenotype [52]. The self-assembly

process has also demonst rated the ability to present bioactive epitopes to cells.

For example, the nanoﬁber scaffold containing the neurite-promoting laminin

epitope isolucinelysinevalinealaninevaline (IKVAV) pentapeptide selec-

tively induced rapid differentiation of neural progenitor cells into neurons [99].

NANOSTRUCTURES FOR TISSUE ENGINEERING/REGENERATIVE MEDICINE 385