Challenges in Soft Tissue EngineeringEser Yuksel,M.D.,1̂ 3 JoshuaChoo,4Matthew Wettergreen,5

andMichael Liebschner, Ph.D.5

ABSTRACT

Soft tissue engineering strategies targeting restoration of volume loss have

inherent critical challenges as they relate to the problem of restoration of defects

with a high volume to surface ratio. We outline the problems associated with the

limitations of translational applications regarding soft tissue engineering strategies

as follows: cell survival, mechanical challenges: macroenvironment (scaffold

collapse and on-the-shelf availability), compositional considerations: microenvir-

onment, inducing malignant behavior, cell migration, and cell exhaustion. These

are discussed with our alternative suggestions for solutions.

KEYWORDS: Soft tissue engineering, tissue engineering, high volume recon-

struction, matrix gradient, multilayered matrix, matrix clearance, breast tissue

engineering

The field of tissue engineering emerged as a

distinct entity in the late 1980s. Prior to this time,

the term ‘‘tissue engineering’’ was being mentioned

in a few papers and discussions, but its precise

definitions remained nebulous. The main impetus

for tissue engineering arose from the challenges

faced in the medical and surgical fields.

Before the concept of tissue engineering

emerged, strategies at addressing tissue losses re-

mained the domain of the surgeon. The removal or

loss of tissue related to disease or trauma often

resulted in functional losses, disfigurement, and, in

cases in which the loss of tissue was incompatible

with life, death. Options were limited to autologous

transfer of the patient’s own tissue, organ trans-

plantation, or the use of artificial prostheses and

implants. All these options continue to have critical

intrinsic limitations. Autologous transfer of similar

tissue is limited or impossible if the tissue loss too

great, as in the case of severe burns, or if the tissue

has specialized metabolic functions. As a result, the

utility of surgical reconstruction using autologous

tissue is usually greatest when the aim is to replace

volume or structural deficits, not metabolic deficits.

Autologous tissue transfers, even when indicated,

introduce the additional problem of donor site

morbidity. Organ tissue transplantation overcomes

some of these limitations; it has been successful in

replacing tissue with vital metabolic functions, such

as liver and kidneys. However, it, too, has serious

Tissue Repair, Regeneration, and Engineering in Plastic Surgery; Guest Editors, C. Randall Harrell, M.D. and Eser Yuksel, M.D. Seminars inPlastic Surgery, Volume 19, Number 3, 2005. Address for correspondence and reprint requests: Eser Yuksel, M.D., 1709 Dryden, Suite 2260,Baylor College of Medicine, Department of Surgery, Division of Plastic Surgery, Houston, TX 77030. 1Methodist Hospital, 2St. Luke’sEpiscopal Hospital, 3Texas Children’s Hospital, 4Baylor College of Medicine, 5Rice University, Department of Bioengineering, Houston,Texas.Copyright # 2005 by Thieme Medical Publishers, Inc., 333 Seventh Avenue, New York, NY 10001, USA. Tel: +1(212) 584-4662.1535-2188,p;2005,19,03,261,270,ftx,en;sps00176x.

261

intrinsic limitations, including the limited availabil-

ity of organs for transplantation and lifelong risk of

immunologic complications. Finally, prosthetic de-

vices and artificial implants, although becoming

increasingly sophisticated, are still limited in their

use to addressing volume or structural deficits

accompanying tissue loss. Furthermore, implant

distortion and rejection secondary to host immune

reactions remain a major problem.

Tissue engineering emerged from the back-

ground of these surgical problems. The concept

represented a major paradigm shift: manipulation

of living cells and their extracellular products in the

development of biological substitutes for replace-

ments as opposed to the use of inert implants or

tissue or organ transfer. Several meetings and sym-

posia, most notably Granlibakken in 1988 (National

Science Foundation), clarified the goals and defi-

nitions of this nascent field. Tissue engineering

was defined as ‘‘the application of principles and

methods of engineering and life sciences toward

fundamental understanding of structure-function

relationships in normal and pathological mamma-

lian tissues and the development of biological

substitutes to restore, maintain, or improve tissue

function.’’1

The landmark article published in Science in

1993 by Langer and Vacanti has since become

the seminal work in the field of tissue engineering.

This paper identified three main strategies for

the development of biological substitutes: the use

of (1) isolated cells or cell substitutes, (2) tissue-

inducing substances, or (3) cells placed on or within

matrices.2

TISSUE ENGINEERING AND VOLUME

RECONSTRUCTION, CHALLENGES

AND SOLUTIONS

All three strategies mentioned by Langer and

Vacanti remain the primary methods employed in

tissue engineering for the purposes of volume re-

storation. The challenges of tissue engineering as

they relate to volume restoration are within the

scope of the current discussion. Currently, the

critical challenges in tissue engineering as they

relate to the problem of volume restoration stem

from the high volume to surface ratio of most

volume defects.

Cell Survival

The high volume to surface ratio severely limits

long-term cell survival because of problem of ad-

equate oxygenation and nutrition (i.e., vascular

access). The goal of increasing the in vivo survival

of engineered cells and allowing organ patterning

(i.e., matching the three-dimensional structural

topography of the lost tissue) has therefore been

one of the primary challenges in tissue engineering

thus far.

PREFABRICATION OF A VASCULAR SUPPLY

One approach to this problem of long-term survival

of engineered tissue implants has been what is

referred to as vascular network engineering: the

use of computer and computed tomography–aided

tissue engineering in constructing a three-dimen-

sional scaffold to direct the growth of the vascular

tree that will support the metabolic demands of

engineered tissue. This has been attempted both ex

vivo and in vivo. Tissue culture systems seeded with

endothelial cells have been used, with limited suc-

cess, in constructing vascular networks in vitro3. In

general, however, it appears to be more promising to

introduce progenitor vascular cells in vivo or rely on

modulating influences within the host organism to

achieve neovascularization in vivo. In this regard,

prevascular components such as stem cell–derived

vascular cells or endothelial progenitor cells have

been introduced in vivo to induce vasculogenesis in

a manner resembling the embryological process in

which ‘‘hemangioblasts’’ differentiate into blood

cells.4 Most attempts have also involved the use of

growth factors such as vascular endothelial growth

factor, and such attempts have shown some success

in creating new vasculature in animal models.5

262 SEMINARS IN PLASTIC SURGERY/VOLUME 19, NUMBER 3 2005

Specific homing, vascular differentiation,

and shape of the resulting vascular tree have yet to

be controlled or studied. The effects of wound

environment (i.e., inflammation, fibrosis, and ex-

tracellular matrix composition) on vasculogenesis

need to be identified and optimized in a differential

manner. In short, the concept of inducing a vascular

support for engineered tissue has a long way to go

before it can be used for clinical applications in

tissue engineering.

LAMINAR TISSUE CONSTRUCTS

One approach we propose to the problem of the

volume/surface area restrictions in the implantation

of engineered tissue is the concept of laminar

construction of engineered tissue. Theoretically,

this would require not the engineering of a complex

vascular tree but the prefabrication of laminar tissue

allowing shorter diffusion distances for oxygen and

nutrients. Vascular supply could be established by

various means, for example, through a comb-shaped

vascular bed using silicon spacers that would sup-

port the laminar insertion of engineered tissue grafts

or progenitor cells. We have used the concept of a

negative template for tissue grafts for the prefabri-

cation of the complex-shaped three-dimensional

structures (Fig. 1).

NONCELLULAR IMPLANTS TO ACHIEVE TISSUE

GENERATION WITH ACCOMPANYING

NEOVASCULARITY: CHALLENGES AND

POSSIBLE SOLUTIONS

A third way to address the problem of volume

restoration is the in vivo insertion of prefabricated

noncellular constructs. The main advantage of this

method is that it activates progenitor cells from the

Figure 1 Simplified schematic presentation of layered tissue and silicon spacers for staged tissue ingrowth with short

distance parameters. (A) Silicon spacer and biodegradable matrix are placed. (B) Regenerated tissue replaces the matrix

(short distance to fill). (C) One tooth unit of silicon spacer is removed. (D) Tissue fills the removed tooth-unit area. (E)

Second tooth unit of silicon spacer is removed. (F) Tissue fills the second tooth-unit area. (G) Tissue tooth unit of silicon

spacer is removed. (H) Tissue growth is achieved within the targeted high volume/surface area.

CHALLENGES IN TISSUE ENGINEERING/YUKSEL ET AL 263

recipient bed, thus inducing soft tissue regeneration

in vivo. Although achieving vascular support is still

a challenge with this method, the gradual induction

of soft tissue replacement allows more time for the

in vivo generation of an adequate vascular supply.

Mechanical Challenges and Structural

Considerations: Macroenvironment

SCAFFOLD COLLAPSE AND ON-THE-SHELF

AVAILABILITY

The concept of using noncellular constructs has its

own challenges that must be addressed before it can

be applied to the restoration of volume defects with

soft tissue. The first challenge is the mechanical

design parameters that the construct must satisfy.

The scaffold would have to match the contours of

the removed tissue and provide initial mechanical

support to prevent collapse of the volume defect

while being able to degrade completely and be

replaced with soft tissue. The degradation proper-

ties of the scaffold would have to be optimized

such that the kinetics match the de novo generation

of soft tissue and the by-products do not harm

the surrounding tissue or hamper de novo tissue

generation.

We have worked on developing such a scaf-

fold according to these design parameters. With the

first design parameter in mind, we have devised a

complex scaffold with three components. These are

a soft fibrillar exterior region that interfaces with the

native tissue and conforms to topographical varia-

tions in the volume defect (facilitates on-the-shelf

availability), a rigid slowly degrading cage portion

that supports load and maintains the geometric

volume, and a highly porous central region housed

within the rigid cage that contains the drug delivery

vehicle and allows tissue ingrowth (Fig. 2).

In our construct, we have used an outer

component composed of a thin layer of fibrillar

matrix that is designed to aid in clotting and

reducing stress profiles secondary to material mis-

match.6 The rigid shell serves multiple functions as

a porous conduit supporting drug delivery and tissue

migration into the architecture and as a mechanical

scaffold preventing tissue collapse.7 This layer can

be fabricated from several slowly degrading polymer

mixes such as poly(lactic-co-glycolic acid) (PLGA).

Additional support will be added to the 2- to 3-mm

layer by patterning the layer with a honeycomb

architecture taken from cellular solids, which are

architectures that optimize strength based on a

specific material volume (Fig. 2).8 The honeycomb

is a mechanically optimized structure that provides

excellent in-plane compressive strength.9

Compositional Considerations:

Microenvironment

Much attention must also be given to the micro-

environment, that is, the environment ‘‘seen’’ by

the cells, which encompasses both structural

elements (i.e., porosity) and chemical makeup (i.e.,

extracellular matrix composition and cell adhesive-

ness characteristics). These elements would facili-

tate the adherence and infiltration of the newly

generated tissue.

The patterned microenvironment of the con-

struct involves many considerations. It must provide

a degradable porous conduit for tissue infiltration,

provide a substrate for cell adherence and migration,

mimic the optimal extracellular environment for the

engineered tissue, and house the delivery vehicles

for any desired bioactive compounds.

The materials that have been used for the

purpose of engineering tissues in vitro are plentiful

and full into two main categories, biological mate-

rials and synthetic materials. Biological materials

include collagen and fibrin matrices and contain an

abundance of biological cell signaling domains,

some of which are well characterized but most of

which are uncharacterized. However, it is clear

that extracellular matrix elements are essential par-

ticipants in cell signaling, cell proliferation, cell

migration, and cell adhesion. The use of biological

materials, although necessary, therefore introduces

additional variables in the bioactivity of the scaffold.

The use of biomaterials may thus result in poorly

264 SEMINARS IN PLASTIC SURGERY/VOLUME 19, NUMBER 3 2005

controlled communication between the scaffold

and the cells that leads to an unregulated result.

Furthermore, the inherent structural characteristics

of such biological scaffolds are not easily modified,

thus limiting control over which physical cues are

received by the cell.

Synthetic scaffolds are typically made from

biologically inert materials, which reduces the

Figure 2 Principles of our scaffold design. (A) Volume defect following partial mastectomy. (B) Placement of scaffold/

delivery vehicle in defect. (C) De novo adipose tissue. (D & F) 3-D and 2-D cut-away view showing the design parameters of

the proposed scaffold construct: (i) 4–5 mm fibrillar layer to fit the contours of the volume defect; (ii) rigid slow degrading

outer shell with honeycomb architecture to prevent structural collapse during tissue ingrowth; and (iii) a fast degrading,

highly porous core containing the microsphere-based PPARg ligand delivery vehicle. (E) Magnification of the central core

showing a microsphere-filled pore. (G) Deformation of the outer fibrillar layer in response to compressive forces allows a

prefabricated scaffold to adjust to a patient’s particular volume defect without the need for extensive tailoring.

CHALLENGES IN TISSUE ENGINEERING/YUKSEL ET AL 265

unpredictability of physical and biological interac-

tions between the cells and biological materials.

Furthermore, the use of synthetic polymers permits

high-resolution control over structural properties

such as density, porosity, and compliance. However,

these materials do not mimic the natural extracel-

lular milieu of the desired tissue as well as natural

materials, thus reducing the stimulatory potential of

the extracellular environment necessary for tissue

engineering aims.

The ability to create synthetic materials that

can be covalently modified with biological materials

has introduced a third category of hybrid materials,

containing both synthetic and natural compounds.

For example, polyethylene glycol (PEG) acrylates,

covalently modified with the cell adhesive peptide

arginine-glycine-aspartic acid (RGD) or the extrac-

ellular matrix component heparan sulfate, have

been successfully incorporated within the scaffolds

to facilitate cell attachment and to allow spatial

sequestration of heparan-binding growth factors.

Several successful attempts at such hybrid

materials exist in the literature. Luo and Shoichet

have created agarose gels with an adhesive fibro-

nectin fragment to facilitate and direct cell migra-

tion during neurite outgrowth.10 Kapur and

Shoichet have designed similar scaffolds that have

incorporated onto a synthetic backbone an immo-

bilized concentration gradient of nerve growth

factor.11 Hubbell has used a PEG backbone that

is cross-linked with proteolytically sensitive oligo-

peptides, allowing cells mobility in the PEG gel

through their own naturally secreted proteases.12

These studies have demonstrated the design

concept of using synthetic polymers, for which the

mechanical and structural properties can be manip-

ulated, and transforming them into biologically

responsive scaffolds for cellular remodeling and

tissue engineering. Hybrid scaffolds such as these

therefore combine the benefits of both synthetic

and biological materials. They allow control over

the structural properties of the scaffold at the

microarchitectural level and permit the addition of

biofunctionality to the scaffold through the incor-

poration of biological materials, which can be

several and include not only extracellular matrix

components but also any other desired biologically

active components, including protein fragments,

growth factors, or even biologically active peptide

sequences. Furthermore, a precise, predesigned spa-

tiotemporal distribution of these factors is possible

within a specifically predesigned microenvironment.

However, our previous methods have not

been able to control the spatiotemporal distribution

of such bioactive compounds. Furthermore, we have

not investigated the necessity of vasculogenic com-

pounds and proteases and extracellular adhesive

components to facilitate the increased demands on

cell proliferation, migration, and vasculogenesis that

would be required in larger volume defects.

Inducing Malignant Behavior

The applied strategies of tissue engineering mimic

the pathways that cancer cells naturally utilize, such

as cell proliferation, cell migration, angiogenesis,

and matrix turnover. These components of cell

induction should be meticulously identified in tissue

regeneration strategies, especially in postcancer re-

construction. Two behavioral models that do not

follow or parallel cancer initiation or progression are

cell differentiation and preventing inflammation.

Therefore, anti-inflammatory and cell differentia-

tive effects of utilized bioactive compounds should

balance or overweigh the aggressive tissue growth-

invasion vector.

In our research, we have use synthetic drug

delivery vehicles such as PGLA-based microspheres

to achieve local long-term delivery of bioactive

compounds such as insulin-like growth factor 1

(IGF-1). We have developed scaffolds with a

porous inner region designed to function as

the drug delivery component. A fast-degrading

50:50 mix of PLGA with 90% porosity was

designed. To offset the reduction in apparent prop-

erties related to the material volume, a strength-

optimized architecture taken from a previous study

evaluating different architectures resulting in varied

apparent properties was used.13 The pore size of

266 SEMINARS IN PLASTIC SURGERY/VOLUME 19, NUMBER 3 2005

polymer was 500 to 800 mm, a size that has been

shown to support tissue invasion in PLGA porous

scaffolds in a rat model.14

We have demonstrated total soft tissue re-

placement of our biodegradable scaffold systems in

2.5-cm defects in a Sprague-Dawley rat model

when combined with adipogenic factors such as

insulin and IGF-1.15 Furthermore, we have shown

that augmentation of adipofascial flaps and free fat

grafts with insulin and IGF-1 delivery vehicles

Figure 3 Findings from our previous work; please refer to the references. (A) Survival rate of fat grafts are improved

110% in insulinþ IGF-1 group (right) compared to control group (left) 110. (B) Volume of adiposo-fascial flap is doubled in

insulinþ IGF-1 group (left upper) compared to control group (right upper) in rat inguinal flapmodel. Note the ectopic de novo

fat islands in insulinþ IGF-1 group. (C) De novo fat islands over abdominal muscular fascia in insulinþ IGF-1 group (center,

SEM of the control group [upper left] versus SEM of the insulinþ IGF-1 group [upper right]).

CHALLENGES IN TISSUE ENGINEERING/YUKSEL ET AL 267

resulted in statistically significant increases in soft

tissue regeneration (Fig. 3), which in both cases was

secondary to adipocyte proliferation and differen-

tiation and not hypertrophy of existing adipo-

cytes.13,15,16 Considering the possible risks with

IGF-1 of inducing malignant behavior (particularly

in the reconstruction of defects of cancer resection),

we have ceased to utilize these agents for postcancer

reconstruction modalities. We have also turned our

attention to the use of biodegradable scaffolds

containing peroxisome proliferator-activated recep-

tor g (PPARg) ligands as our long-release compo-

nents in a biodegradable microsphere system.

PPARg receptors play a pivotal role in the terminal

adipocytic differentiation of precursor cells. Like

insulin and IGF-1, PPARg ligands have been

shown to stimulate adipocyte proliferation and

differentiation; unlike insulin and IGF-1, PPARgligands do not have the same tumorigenic effect

and indeed have demonstrated an antineoplastic

effect on breast cancer cell lines, as mentioned

previously.17 These effects are possibly due to the

differentiative and anti-inflammatory properties of

PPARg ligands.

Cell Migration

This ability to control precisely the spatiotemporal

distribution of the biological factors is of tremen-

dous import in tissue engineering and will allow

the creation of pull vectors and gradients guiding

cell migration. Matrix modification is an important

component of this process as chemotactic signals in

Figure 3 (continued) (D) Insertion of a 2.5 cm porous PLGA implant over the latissimus dorsi region in Sprague-Dawley rat

model (left); SEM of the PLGA scaffold with microspheres (right). (E) Adipose tissue totally replacing the PLGA implant

over the latissimus dorsi region in Sprague-Dawley rat model at 12 weeks (left); light microscopy of the fibro-fatty tissue

over the latissimus dorsi muscle at 12 weeks (right).

268 SEMINARS IN PLASTIC SURGERY/VOLUME 19, NUMBER 3 2005

the matrix are required to stimulate inward move-

ment of cells. In vivo design of matrix composition

should be aimed at targeting the specialized cells of

the target tissue (e.g., preadipocytes for soft tissue

regeneration, endothelial cells or hemangioblasts for

vasculogenesis).

In addition, micropatterning of biological

factors, including synthetic backbones modified

with alternating adhesive molecules, may make it

possible for cells to ‘‘walk’’ on a scaffold by dynamic

adhesion/contra-adhesion, thus maximizing tissue

infiltration of even large-volume scaffolds.

Micropatterning of scaffolds can also encom-

pass the controlled modification of synthetic back-

bones to create a gradient of pull for the target cell.

This can include bioactive proteins incorporated

into the synthetic backbone, the incorporation of

synthetic drug delivery systems suspended within

the scaffold in a gradient-like fashion, or the crea-

tion of an electrochemical gradient (Fig. 4).

Cell Exhaustion

Regeneration succeeds only to a certain depth.

Several factors contribute to this limitation, includ-

ing increasing resistance to cell migration secondary

to fibrosis, mechanical properties of the scaffold,

and inadequate vascular supply. These limitations

have been the main difficulty in translating the

results of rat studies to higher volume defects.

We can utilize the strategy of the cancer cell to

overcome this problem, such as releasing matrix

degradation enzymes to clear the matrix wall. One

Figure 4 Patterning and creating gradient within the matrix. (A) Proximo-distal and centripedal gradient (in a nerve matrix

model). (B) ‘‘Repellent-attractant’’ alternating matrix patterning for cell-walk, by modifying/coating with adhesive and

counter adhesive ECM proteins. (C) Surface patterning with alternating ECM proteins and creating a proximo-distal and

centripedal gradient distribution of delivered bioactive molecules—according to their charge characteristics, utilizing a

magnetic field in the same direction.

CHALLENGES IN TISSUE ENGINEERING/YUKSEL ET AL 269

possible solution is the delivery of tissue-type plas-

minogen activator (tPA) or other matrix degrada-

tion enzymes to facilitate the degradation of the

extracellular environment necessary for cell migra-

tion and vasculogenesis. Theoretically, one can

control the spatiotemporal distribution of such

enzymatic activity along an increasing gradient of

proteolytic activity normal to the noncellular scaf-

fold. This can be done through selective modifica-

tion of the synthetic backbone in a stratified manner

to create shells of increasing tPA activity. A supra-

physiologic concentration of tPA may be necessary

to maximize cell migration past a critical distance.

VASCULOGENESIS

The vascularization of engineered tissues in many

cases does not keep up with the ingrowth of cells.

Nutrient and oxygen supply are not sufficient,

which ultimately leads to the death of the invading

cells. The enhancement of the angiogenic capabil-

ities of engineered tissues therefore represents a

major challenge in the field of tissue engineering.

The immobilization of angiogenic growth factors in

the same manner as tPA may be useful for enhanc-

ing angiogenesis.

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