Plastic Surgery Volume 1 Principles 4th

226
CHAPTER 15 • Skin graft
Full-thickness donor sites in the head and neck are pre- and
postauricular regions (the first is generally thicker), nasolabial
crease, supraclavicular region, eyelids, and neck.
Other common regions include the inguinal crease, which
is often used to cover large defects. In this case it is important
to harvest laterally and away from potentially hair-bearing
regions of the pubis. This area generally heals well and is
hidden (see Fig. 15.11).
In nipple–areola reconstruction after mastectomy, a
full-thickness graft from the contralateral region is often
taken in combination with a full-thickness graft from the
groin.
Donor sites from full-thickness skin grafts are generally
closed primarily. Particular attention should be taken when
supraclavicular donor sites are used, as this region is prone
to develop hypertrophic scars. Sometimes large donor sites
that cannot heal primarily are used for covering extensive
defects, for example, in the face or joints. In this case, an STSG
can be used to cover the donor site.
Donor site dressing
Topical gauze soaked with diluted epinephrine solution is
useful to stop bleeding at the donor site and can be left until
the operation is ended, adding an analgesic effect. The donor
site of an STSG generally heals (re-epithelializes) in 7–21
days depending on the size and depth of the graft taken and
the age of the patient. A myriad of donor site dressings are
available with multiple studies on a variety of products.
Traditionally, fine-meshed gauze, often impregnated with a
petroleum-based product, is placed over the wound and
fixed in place. Cotton gauze is placed over this and removed
a day or two after the operation. The wound heals under
this dressing and the dressing spontaneously comes off
when healed. The advantages of this type of dressing system
are its simplicity, low cost, and minimal wound care require­
ment. The work of Winter in the early 1960s51 showed that a
moist wound heals faster. As a consequence many have
advocated a simple semiocclusive, adhesive, semipermeable
polyurethane dressing, although it promotes serum accumu­
lation between the wound and the dressing that should be
evacuated by a drain or a syringe. This process can be labor-
intensive for a busy inpatient service or difficult to manage
on an outpatient basis. In addition, when an infection devel­
ops, it can rapidly spread to the entire donor site area with
this dressing.
A large number of other dressings, including silver-based,
absorptive, and biological, have been studied. In most of the
cases the donor site heals spontaneously, thus simple impreg­
nated gauze is still the gold standard. When some complica­
tions occur, the harvest site can deepen and become a
full-thickness defect, mainly due to infections in the elderly,
infants, or critically ill patients. As a consequence, excision
and grafting of the donor site may be required.
Skin graft storage
Skin grafts can be stored on a moist gauze at 4°C for up to 2
weeks, although the viability decreases over time. Experimen­
tally, storage can be extended using cell culture media. For
degloving injuries, the skin can be defatted acutely and reap­
plied as a full-thickness graft.
replacement of tissue from a similar or adjacent site will give
the best color match. In the face, this is often the most critical
aesthetic area and choosing donor sites such as supraclavicu­
lar, posterior auricular, upper eyelid or scalp skin grafts can
often lead to an excellent color match. For darker areas of the
skin, such as the nipple–areola complex, grafts from the
contralateral areola or genitalia have been used. More com­
monly today, skin grafts in this area can be tattooed with
vegetable dyes to give an excellent color match.
Skin texture is most commonly an issue when dealing with
glabrous skin (palms and soles of feet). In this case, placing a
nonglabrous skin graft to cover these areas can result in a very
unnatural look, often with significant color match differences.
Glabrous skin grafts are obviously in short supply but can be
harvested from the hypothenar eminence. Skin adnexal glands
are difficult to replace but can sometimes be successfully
transferred with full-thickness grafts.
Donor site considerations
The obligatory scarring or discoloration associated with donor
sites must be considered when taking a skin graft (Fig. 15.14).
For large surface area burns, the surgeon must use whatever
donor site is available to close the wound; in contrast, for
smaller grafts, a thorough discussion of the donor site possi­
bilities with the potential risks and benefits allows for the best
skin match with the least iatrogenic damages.
Common donor sites include the thigh, trunk, and
buttocks, regions frequently covered by clothing (see
Fig. 15.4).
When defects of the face need to be covered, full-
thickness skin grafts are often preferred. The donor area for a
graft on the face should be preferably chosen from among the
scalp, neck, and supraclavicular area to obtain the best
color match. Important aspects are the thickness, color, and
texture of the graft, which should closely match those of the
recipient site.
Eyelid skin, which is thin with a few glandular structures,
is generally best replaced by eyelid skin. Thick, highly glan­
dular nasal skin can be replaced by skin from the nasolabial
folds, supraclavicular area, or anterior auricular area.
In trauma, avulsed (degloved) skin can often be prepared
by extensive defatting and regrafted primarily50 or stored and
used later.
Fig. 15.14  Donor site three years after split-thickness skin graft taken with a
dermatome at the forearm. Note slight hypopigmentation.

227
Future
extracellular matrix becomes mechanically resistant. This
increase in mechanical resistance allows mechanical interac­
tion with fibroblasts that develop fibers called alpha-SMA.
The development of alpha-SMA coincides with the differen­
tiation of fibroblasts into myofibroblasts that induce wound
contraction.
Instability
Shear forces are a major cause of skin graft failure that can
disrupt the nascent fragile blood vessel connections. Later,
thin skin grafts have less collagen content and are prone to
delamination due to shear forces.
Cosmetic issues
Unpleasant cobblestone-like patterns of meshed STSGs are
much more prominent than nonmeshed or full-thickness
grafts. Color differences are a common problem if donor sites
could not be optimally chosen. If excess skin surrounds the
grafted area, the skin graft can be excised in multiple steps
until primary closure can be reached.
Donor site
Infection at donor sites can occur from bacterial contamina­
tion during the operation or in the postoperative period.
These infections are treated with topical antimicrobial agents,
including silver dressings. The delay in healing of donor sites
due to infection can lead to hypertrophic scar formation.
Hypertrophic scar or keloid formation can also result from
deep donor sites or in patients with a propensity for scarring.
Itching is a common reaction to donor sites as well as hyper­
sensitivity to changes in temperature.
Future
Dermal substitutes
Burke et al. first described a collagen–glycosaminoglycan
scaffold to treat dermal defects in the early 1980s.54 Dermal
substitutes (Table 15.3) are widely used in large burn injuries,
when skin lesions are so extensive and unaffected donor skin
areas are often limited. Cadaver skin (allograft) or dermal
substitutes can be used initially to cover large defects. Dermal
substitutes can also be used to improve functional and aes­
thetic outcome. Dermal regeneration is very limited in human
skin and full-thickness skin defects heal by secondary inten­
tion with scar tissue formation. The difference between scar
tissue and healthy skin besides the aesthetic appearance is the
viscoelasticity and therefore stability. Scar tissue shows histo­
logical parallel-oriented collagen fibers, while healthy dermis
consists of randomly oriented collagen fibers. In some patients
and for restricted donor sites, keratinocyte cultures or very
thin STSGs can be attempted to close full-thickness defects.
The resulting skin is usually fragile and dry, leading to
reopening and ulcers through normal mechanical friction.
This poses a particular problem if functional structures under­
lie the new thin skin. The introduction of the first commercially
available skin substitute, Integra®, is a bovine-derived colla­
gen type I cross-linked matrix with glycosaminoglycans that
is similar to the dermal structure and covered by a silicone
Complications
Hematoma
Any liquid between the wound and skin graft can impair skin
graft take. Bleeding represents one of the most important
complications after excision and grafting, with up to
100–200 mL of blood for every 1% of body surface area that
is excised, particularly from the scalp.52 The surgeon must be
certain that bleeding has stopped prior to dressing the wound.
Suture ligation or cautery can be used to control larger bleed­
ing vessels; oozing can be controlled with pressure and/or
pharmacological methods such as topical thrombin, epineph­
rine, or fibrin glue. To reduce bleeding during excision the
area can be primarily injected with epinephrine diluted in
saline (tumescent technique: see Fig. 15.5).53 As discussed
above, incisions of the graft should be done to allow fluid
evacuation through a compressive or suction dressing.
Seroma
Serum imbibition is essential for early skin graft survival.
Excessive serum, such as a seroma, will prevent or delay skin
graft take. Seromas are better tolerated than hematomas, and
adequate fenestrations (pie crusting) can prevent this problem.
Infection
When skin graft infections occur, pus often accumulates
beneath the skin graft and can rapidly spread. If an infection
is found early, prompt incision and drainage of the fluid
beneath the graft can often salvage some or all of the skin
graft. A large number of dressings have been developed that
carry a variety of topical antimicrobial agents. Silver nitrate,
mafenide acetate, and silver ion dressings are commonly
used. Bacteria seem not to develop resistance easily to silver
products, making these products desirable.
A contaminated wound will not heal and will reject a graft.
Some microorganisms, such as Pseudomonas spp., can con­
taminate the wound and cause nonpurulent infections. Even
systemic or nonskin-localized infections have been proven to
delay or prevent skin graft taking.
Nontake
Unfavorable conditions such as malnutrition, vasculitis,
malignant diseases, steroids, and chemotherapeutic medica­
tions have all been shown to cause or accelerate nontake of
the graft.
Wound contraction
Wounds covered with skin grafts can still undergo wound
contraction leading to a scar contracture. After the acute
inflammation phase the contraction starts. Fibroblasts from
surrounding tissues and from blood circulation start to be
activated and repopulate the wound between the graft and
the wound bed. Cells at the interface need to build new
collagen-rich tissue to replace fibrin and anchor the graft to
the wound bed. While fibrin is crucial for early cell migration
in the wound bed, collagen deposition allows wound matura­
tion. As collagen is deposited, cross-linking occurs so that the

228
CHAPTER 15 • Skin graft
dressings will be removed and a well-vascularized wound
bed is found underneath that is ready for final skin recon­
struction by autologous grafts. Infection of the acellular
dermis demands careful wound dressing and infection
control. Another drawback for human- and porcine-derived
dermis is possible viral infection transmission.55
Cell cultures
Epithelial cell culture autografts were first introduced by
Rheinwald and Green in 1975 and pose a milestone in skin
regeneration.56 In patients with extensive burns, donor sites
are often limited. Cultured epithelial autografts (CEA) are
keratinocytes harvested from a small biopsy of the same
patient that are then expanded manifold in the laboratory
(Fig. 15.16).
The time needed to expand keratinocyte cultures in vitro
for clinical use is dependent upon the delivery method. In
order to obtain sheets of confluent keratinocytes as in normal
epidermis, it may take up to 5 weeks. This time can be short­
ened to less than 2 weeks57 by expanding the cultures on
bioscaffolds that allow cell attachment and proliferation prior
to transplantation. Hyaluronan or collagen scaffolds have
been demonstrated to be very useful in delivering cell-seeded
sheets with approximately 80% confluence up to even multi­
layered epithelial tissue constructs. Cell suspensions have the
advantage of being delivered faster, reducing the in vitro time
to only 5–7 days. Recently, ReCell® (Avita Medical), invented
sheet to recreate temporarily the function of the epidermis.
Integra® can be applied to a vascularized wound bed followed
by blood vessels and cell growth into the collagen matrix to
create a vascularized neodermis after 2–3 weeks. During the
avascular period, the dermal graft is very sensitive to infec­
tion, one of the main disadvantages of the two-step strategy.
Once revascularization is accomplished, the silicone layer
can be removed and the epidermal replacement can be granted
by either keratinocyte cultures or a thin STSG (Fig. 15.15). To
improve dermal graft integration, surgeons have successfully
seeded dermal substitutes with keratinocytes.23 Newer clini­
cal studies show that small areas of exposed tendon or bone
without remaining peritendineum or periosteum can be suc­
cessfully grafted with Integra®.
After the invention of Integra®, a number of other human-
and animal-derived dermal substitutes are now commercially
available. Another synthetically engineered dermal substitute
is MatriDerm® (MedSkin Solutions Dr. Suwelack), a bovine-
derived collagen matrix with noncross-linked collagen fibers
of types I, III, and V matrix coated with elastin fibers. Synthetic
dermal substitutes such as Integra® and MatriDerm® are
permanent dermal regenerative templates that will be ulti­
mately covered by an epidermal autologous graft. For tempo­
ral coverage of large wounds, dermal allografts or xenografts
are available (see Table 15.3). Allografts and xenografts will
be rejected by the recipient and may adhere in patients with
large burn injuries for several weeks because of the post-
traumatic immunosuppressed stage. After 2–3 weeks dermal
Table 15.3  Permanent and temporary dermal and epidermal skin substitutes
Product
Tissue – cells
Manufacturing availability
Origin
Epicel® (Genzyme, MA)
Epidermis – cultured epidermal autograft
(CEA) sheets
Tissue cultures expanded in
the laboratory over several
weeks
Autologous and
xenogeneic
(residual amounts
of murine cells)
ReCell® (Avita Medical, UK)
Epidermis – autologous epidermal cells
Dermis – fibroblasts; cell suspension,
delivered with a spray
Bedside approach (about 30
minutes required)
Autologous
Cellutome™ Epidermal
Harvesting System (KCI, TX)
Epidermis – autologous epidermal islands
delivered on a dressing
Bedside approach
(about 60 minutes required)
Autologous
Integra® (Integra Lifesciences,
NJ)
Dermis – bovine tendon type I collagen and
glycosaminoglycans on a silicone sheet
On the shelf
Xenogeneic
MatriDerm® (MedSkin Solutions
Dr. Suwelack, Germany)
Dermis – bovine acellular non-crosslinked,
coated with elastin
On the shelf
Xenogeneic
Alloderm® (LifeCell Corporation,
NJ)
Dermis – human acellular lyophilized
cadaver dermis
On the shelf
Allogeneic
Dermagraft® (Organogenesis,
MA)
Dermis – human fibroblasts on polyglycolic–
polylactic acid mesh
On the shelf
Allogeneic
EZ Derm (Mölnlycke Health
Care, Sweden)
Dermis – porcine aldehyde cross-linked
dermal collagen
On the shelf
Xenogeneic
Oasis® Matrix (Smith and
Nephew, TN)
Dermis – porcine acellular small intestine
submucosa
On the shelf
Xenogeneic
Allograft
Composite – cryopreserved cadaveric skin
On the shelf
Allogeneic
Apligraf® (Organogenesis, MA)
Composite – neonatal human fibroblasts in
bovine type I collagen
Neonatal human keratinocytes
On the shelf
Allogeneic/
xenogeneic

229
Future
efficacy and indications, has not been identified yet. Published
studies have reported take rates of CEA on biomaterials
varying from 0 to 80%, raising many questions about their
effectiveness. Boyce et al.59 described bilayered autologous
skin constructs with keratinocytes and fibroblasts on collagen–
glycosaminoglycan matrixes that showed promising results in
burn treatment. Direct comparison of autologous bilayered
skin constructs versus STSGs showed higher initial nontake
rates and regraftment after treatment with the skin construct,
while STSG-treated areas healed spontaneously. Once wound
closure occurred, the outcomes showed similar results in skin
quality at longer time points. The variability may be explained
by heterogeneous wound cohorts, inconsistencies in the time
at which wounds were analyzed, treatment techniques, and
measuring systems.
Bioengineered cultured allogeneic
bilayered constructs
Cell-seeded skin constructs were first described in the early
1980s, when autologous keratinocytes and fibroblasts were
Fig. 15.15  Skin regeneration with dermal substitutes. If large, deep skin defects
have to be covered with thin split-thickness skin grafts, dermal substitutes can be
used to augment the dermal layer and therefore the aesthetic and functional
outcome. (Adapted from )
A
Undamaged epidermis
Undamaged dermis
Silicone membrane
Dermal substitute
B
Blood vessel ingrowth
C
Silicone membrane
removed
D
Meshed skin graft
E
Regenerated skin
Fig. 15.16  Keratinocyte cultures. Epidermal cells can be harvested from a small
biopsy from the patient and expanded manifold in vitro. Once the expansion is
achieved the autologous keratinocytes can be delivered back to the patient.
(Adapted from Autologous Cultured Epidermis, )
Skin biopsy
Skin specimen
Isolation of keratinocytes
Feeder cells
Cell seeding/innoculation
Culture
Sheet formation
Grafting
Injured area
by Fiona Wood et al., introduced an even faster method,
applying noncultured keratinocytes in a one-step procedure.58
Since in deep wounds time is a critical issue, this strategy,
offering accelerated delivery, is promising, although still
poorly characterized. One major concern is whether cells
directly sprayed on the wound surface are able to find attach­
ment sites to engraft and survive. Another bedside approach
for epidermal graft is represented by the Cellutome™ (KCI,
TX), which allows for harvesting epidermal islets down to the
basal layer. The islets are then transferred on the recipient
wound via a dressing and the donor site heals spontaneously
in a few days.
The introduction of CEA in clinical practice dates back
more than 20 years, but the best delivery method, as well as

230
CHAPTER 15 • Skin graft
keratinocytes and fibroblasts obtained from neonatal foreskin.
Fibroblasts are cultured in the collagen matrix where they
proliferate and augment the extracellular matrix with all types
of proteins. Keratinocytes are then added and build up the
epidermal layer. The living skin construct supports skin graft
take in difficult wounds, such as burn wounds, or accelerates
healing in chronic wounds. With a shelf-life of 5 days at room
temperature, Apligraf® has to be applied fresh either as a
temporary dressing or over meshed STSGs.
Future products that are less immunogenic might be useful
for improving skin regeneration, and novel technologies such
as 3D printing will soon make clinically available more
sophisticated synthetic constructs.
seeded into collagen glycosaminoglycan matrixes by centrifu­
gation.60 Further research found neonatal skin cells to be very
efficient in accelerating wound healing. Neonatal foreskin
keratinocytes and fibroblasts were than used in combination
with biological or synthetically engineered scaffolds to stimu­
late wound healing in topically applied allogeneic skin con­
structs. Other than autologous cells, which seem to integrate
into wound tissues, allogeneic constructs rather stimulate
healing by growth factor release and initial production of
extracellular matrix proteins, but are rejected eventually.
Apligraf® (Organogenesis, Canton, MA and Novartis Pharma­
ceuticals, East Hanover, NJ) is a bilayered living skin equiva­
lent composed of type I bovine collagen and allogeneic
6. Tanner JC Jr, Vandeput J, Olley JF. The mesh skin graft. Plast
Reconstr Surg. 1964;34:287–292.
12. Blanpain C, Fuchs E. Epidermal stem cells of the skin. Annu Rev
Cell Dev Biol. 2006;22:339–373. Review of the current knowledge of
epidermal stem cells of the adult skin.
14. Fuchs E. Scratching the surface of skin development. Nature.
2007;445:834–842. Review article giving insight into recent developments
that have focused on epidermal stem cell population to maintain hair
follicle regeneration and re-epithelialization in response to wound healing.
29. Converse JM, Smahel J, Ballantyne DL Jr, et al. Inosculation of
vessels of skin graft and host bed: a fortuitous encounter. Br J Plast
Surg. 1975;28:274–282.
33. O’Ceallaigh S, Herrick SE, Bennett WR, et al. Perivascular cells in a
skin graft are rapidly repopulated by host cells. J Plast Reconstr
Aesthet Surg. 2007;60:864–875.
38. Capla JM, Ceradini DJ, Tepper OM, et al. Skin graft vascularization
involves precisely regulated regression and replacement of
endothelial cells through both angiogenesis and vasculogenesis.
Plast Reconstr Surg. 2006;117:836–844.
56. Rheinwald JG, Green H. Serial cultivation of strains of human
epidermal keratinocytes: the formation of keratinizing colonies
from single cells. Cell. 1975;6:331–343. A milestone article in skin
regeneration, introducing the cultivation of keratinocytes for autologous
cell treatment.
59. Boyce ST, Goretsky MJ, Greenhalgh DG, et al. Comparative
assessment of cultured skin substitutes and native skin
autograft for treatment of full-thickness burns. Ann Surg.
1995;222:743–752. Clinical study comparing an autologous bilayered skin
construct based on keratinocytes, fibroblasts, and a collagen
matrix versus the gold standard therapy: autologous skin graft in burn
patients.
61. Orgill DP. Excision and skin grafting of thermal burns. N Engl J
Med. 2009;360:893–901. Review article on current guidelines of the
treatment of thermal burns.
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230.e1
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vacuum-assisted closure (VAC) as a negative pressure dressing in
lower limb split skin grafting: a clinical trial. J Plast Reconstr Aesthet
Surg. 2008;61:334–337.
48. Korber A, Franckson T, Grabbe S, et al. Vacuum assisted closure
device improves the take of mesh grafts in chronic leg ulcer
patients. Dermatology. 2008;216:250–256.
49. Foster K, Greenhalgh D, Gamelli RL, et al. Efficacy and safety
of a fibrin sealant for adherence of autologous skin grafts to burn
wounds: results of a phase 3 clinical study. J Burn Care Res. 2008;29:
293–303.
50. Meara JG, Guo L, Smith JD, et al. Vacuum-assisted closure
in the treatment of degloving injuries. Ann Plast Surg. 1999;42:
589–594.
51. Winter GD. Formation of a scab and the rate of epithelialization of
superficial wounds in the skin of the young domestic pig. Nature.
1962;193:293–294.
52. Cartotto R, Musgrave MA, Beveridge M, et al. Minimizing blood
loss in burn surgery. J Trauma. 2000;49:1034–1039.

230.e2
CHAPTER 15 • Skin graft
53. Robertson RD, Bond P, Wallace B, et al. The tumescent technique to
significantly reduce blood loss during burn surgery. Burns.
2001;27:835–838.
54. Burke JF, Yannas IV, Quinby WC Jr, et al. Successful use of a
physiologically acceptable artificial skin in the treatment of
extensive burn injury. Ann Surg. 1981;194:413–428.
55. Department of Health and Human Services. Human Tissue Intended
for Transplantation (Final Rule). Fed Regist. 1997;62:40429–40447.
Available online at:: <
briefing/3681b1_02_01.htm>.
56. Rheinwald JG, Green H. Serial cultivation of strains of human
epidermal keratinocytes: the formation of keratinizing colonies
from single cells. Cell. 1975;6:331–343. A milestone article in skin
regeneration, introducing the cultivation of keratinocytes for autologous
cell treatment.
57. Zacchi V, Soranzo C, Cortivo R, et al. In vitro engineering of human
skin-like tissue. J Biomed Mater Res. 1998;40:187–194.
58. Wood FM, Stoner ML, Fowler BV, Fear MW. The use of a non-
cultured autologous cell suspension and Integra dermal
regeneration template to repair full-thickness skin wounds in a
porcine model: a one-step process. Burns. 2007;33:693–700.
59. Boyce ST, Goretsky MJ, Greenhalgh DG, et al. Comparative
assessment of cultured skin substitutes and native skin autograft
for treatment of full-thickness burns. Ann Surg. 1995;222:743–
752. Clinical study comparing an autologous bilayered skin construct
based on keratinocytes, fibroblasts, and a collagen matrix versus the gold
standard therapy: autologous skin graft in burn patients.
60. Orgill DP, Butler CE, Barlow M, et al. Method of skin regeneration
using a collagen-glycosaminoglycan matrix and cultured epithelial
autograft. Vol. 5,716,411. U.S. Patent, Feb 10, 1998.
61. Orgill DP. Excision and skin grafting of thermal burns. N Engl J
Med. 2009;360:893–901. Review article on current guidelines of the
treatment of thermal burns.

16 
Tissue engineering
Andrea J. O’Connor, Diego Marre, Kiryu K. Yap, Daniel E. Heath, and Wayne A. Morrison
SYNOPSIS
■ Tissue engineering, i.e., a growth of three-dimensional (3D) tissues
using cells, scaffolds, and blood supply, calls on the skills and
expertise of cell biologists, engineers, and clinical surgeons.
■ It aims to replicate the embryological process of tissue regeneration
rather than adult healing and, given the complexity of the process,
much of our effort to date may seem simplistic.
■ Nevertheless, much progress has been made across the full spectrum
of the research field. This chapter outlines the historical beginnings of
tissue engineering and discusses the current status, addressing the
three structural components of tissue: cells, matrix, and vascular
supply.
■ Biomaterials science and applications are discussed as well as an
overview of emerging technologies in the field.
■ Much of the work in tissue engineering has been performed ex vivo
and unfortunately many of these experiments do not predict the fate of
cells and matrices in vivo.
■ Models of in vivo tissue engineering are discussed along with the
progress in specific fields.
■ Very few clinical trials have been undertaken, controls or placebo
groups are difficult to incorporate in trial design, and tracking the fate
of implanted cells is difficult in humans.
■ All this needs to be viewed with the perspective that, to date, very few
truly tissue-engineered products have come to market despite huge
expectations and investments.
Introduction
Traditional reconstructive plastic surgery techniques “rob
Peter to pay Paul”. The long-held hope for the ultimate
reconstruction has been allotransplantation but to date drug
morbidity makes the risk–benefit ratio unfavorable for non­
vital organs except in rare circumstances. Tissue engineering
has emerged as a possible new direction where ideally our
own body is stimulated to generate its own replacement
tissue.
Definition of tissue engineering
The definition of tissue engineering quoted by most authors
is that it is “an interdisciplinary field that applies the principles
of engineering and the life sciences toward the development
of biological substitutes that restore, maintain, or improve
tissue function”.1 It stressed the key roles of engineering
techniques, material science, biochemical expertise, and cell
biology. What it did not stress was the role that blood supply
would play if and when such engineered tissues were to be
transplanted into the body. It is to this piece of the tissue-
engineering puzzle that plastic surgeons may hold the key
and why the clinical application of tissue engineering will
become the province of the plastic surgeon.
Regenerative medicine
While tissue engineering has predominantly surgical roots
with research directed by cell biologists and engineers, regen­
erative medicine, a modernized branding for cellular therapy,
was the province of physicians and cell biologists. It has
since expanded to target organ regeneration and there is
now considerable crosstalk between the disciplines. For the
purpose of this chapter, we distinguish tissue engineering as
the pursuit of 3D structural tissue generation.
The classical paradigm of tissue engineering followed Jack
Burke’s artificial dermis skin model of seeding cells on to a
biodegradable matrix scaffold and implanting it into a vascu­
larized bed.2 This would rapidly nourish the graft and, by a
combination of survival of implanted cells and substitution,
the “new skin” would become functional. Clinical needs for
bone and cartilage determined early research directions and
this gave pre-eminence to the engineering materials side of
the tissue-engineering pendulum as structural load, stability
and ideally biodegradability were clearly essential. The spec­
tacular image of the “human ear” in the mouse (Fig. 16.1)
galvanized scientists and the public alike as to the potential
offered by tissue engineering and reinforced the dogma of
expanding cells in culture, seeding them ex vivo into scaffold

232
CHAPTER 16 • Tissue engineering
Embryology
Tissue engineering at its core is an attempt to replicate embry­
ology or fetal healing. Here, new tissues form in response to
certain signals, both biochemical and biomechanical.8 The
tissues are determined or patterned according to a genetically
guided time sequence where growth factors and their chemi­
cal gradients as well as changing physical forces in the matrix
attract, guide, and differentiate primitive cells into functional
clusters. In this process, cells will link with a capillary bed,
lay down their own matrix, and create an environment appro­
priate to their expansion and development. This relatively
new concept of “self-assembly” is important to appreciate
when translating embryogenesis into a model for tissue
engineering. As cells change their phenotype, for example,
from precursor to differentiated or from motile and prolifera­
tive to resting and clustered, their ECM requirements alter
accordingly. Many cells secrete their own matrix suitable to
their needs. This constant change cannot easily be replicated
in the laboratory and it is simplistic to believe we can seed
cells on to one specific matrix and assume this will be appro­
priate for the cells’ needs throughout their development in
vivo, nor is it easily conceivable that “smart surfaces” can be
imprinted on synthetic matrices in layers programmed to be
released according to the changing cells’ needs when at this
point we do not know what these needs are.
Examples from nature
Nature has given us many examples of spontaneous tissue
regeneration or engineering which give clues to how this
might be modeled artificially. The liver has been known since
ancient times to self-regenerate and is captured in the legend
of Prometheus, a Titan, who stole fire from Zeus and gave it
to mortals. As punishment, he was bound to a rock while a
great eagle ate his liver each day, only to have it grow back to
be eaten again.
A salamander regrows its legs, the fetus repairs wounds
with minimal scarring, and humans have the capacity to gain
and lose adiposity rapidly. Ruptured tendons can regenerate
across gaps when their ends are retained within their synovial
sheath where their matrix and cellular environment are main­
tained and axial mechanical force signals are transduced into
biochemical stimulation (Fig. 16.2). Muscles also will regener­
ate across gaps in the same way.
Access the Historical Perspective section, including Fig.
16.3, online at

Components of tissue engineering
Adult wound healing involves deliberate sloughing of
damaged tissue and repair by rapid wound closure, scar, and
contracture. A primary goal of tissue engineering is to achieve
a more scarless regeneration through the use of stem cells,
matrix, and growth factors delivered according to a sequence
and type learned from embryology, as described above.
Simplistically, our tissues can be viewed as a composite of:
(1) cells (both parenchymal and stromal); (2) matrix; and (3)
blood vessels. Each of these components must be considered
matrix and implanting the composite into the body.3 The
journal Tissue Engineering was established in 1996 and many
publications focused on novel synthetic biomaterials attempt­
ing to replicate pore size and connectivity of nature’s extracel­
lular matrix (ECM) using sophisticated machinery and
mathematical modeling. Toxicity, structural integrity, biode­
gradability, surface modifications, cell survival, adherence,
migration, and proliferation could all be tested in ex vivo
models. Bioreactors could greatly expand cell numbers and
mimic in vivo physical environments to allow more sophisti­
cated ex vivo experimentation.
The ex vivo paradigm stumbled when attempts were made
to translate advances into the living animal. Biomaterials
proved toxic to cells in some cases and acted as foreign bodies.4
Their dynamic structural behavior was very different to ex vivo
testing and cells found life hostile in the new world. Despite the
obvious need for multidisciplinary cooperation, most early
research was directed in isolation and this has led to delayed
progress in tissue engineering. The engineer’s mindset is
focused on scaffold structure, tensile strength, biocompatibil­
ity, and surface modification while the cell biologist’s is on cell
identification, expansion, and differentiation. Scaffold model­
ing was based on sophisticated replication of human ECMs
with respect to pore size, but evaluation was largely done in
vitro. Both teams are laboratory-based, observing cell responses
outside the living animal, and had limited access to a practicing
clinician’s observations, especially plastic surgeons, with
respect to the behavior of foreign bodies in vivo. Furthermore,
surgeons know from bitter experience with skin grafting that
cells seeded on to an implanted scaffold ex vivo and then
implanted are unlikely to connect to a blood supply in sufficient
time to avoid ischemic death. Vascularization is vital to cell
survival and is a key to any successful multidisciplinary part­
nership towards clinical translation of tissue engineering.
Unfortunately, early tissue engineers were beguiled by select­
ing tissues that were very thin, such as skin, which could
rapidly connect to a blood supply or tissue that did not need a
blood supply, namely cartilage. Once 3D soft tissues were
attempted, it became evident that cell survival was poor and
that blood supply was paramount.5–7
Fig. 16.1  “Human ear in the mouse”: cultured bovine cartilage cells seeded on a
PLGA scaffold in the shape of a human ear and implanted in the back of a mouse.
(Reproduced from Cao Y, Vacanti J, Paige K, et al. Transplantation of chondrocytes
utilizing a polymer-cell construct to produce tissue-engineered cartilage in the shape
of a human ear. Plast Reconstr Surg. 1997;100:297–302.)

232.e1
Historical perspective
Fig. 16.3  (A) Thumb stump with distraction device. (B) X-ray of bone distraction. (C) New bone filling the gap.
A
B
C
Historical perspective
The term “tissue engineering” was coined at a meeting spon­
sored by the National Science Foundation in 1987. Implied in
the definition of tissue engineering is the use of living cells,
biomolecules, and/or biomaterials.9 One of the earliest human
applications of the concepts of tissue engineering is in oral
surgery with what is termed “guided bone regeneration”
(GBR) or guided tissue regeneration (GTR).10 This surgical
procedure utilizes barrier membranes to direct growth of
bone and soft tissue and is predominantly applied in the oral
cavity to support new bone growth on alveolar ridges to allow
stable insertion of implants.11 The membrane which is inserted
under the mucoperiosteum creates a dedicated space that
collects the raw ingredients of osteogenesis – periosteal and
bone-derived osteoblasts, vascular cells, tissue fluid, and
matrix secretions appropriate for the biochemical signaling of
bone formation. The rigid surface of a membrane induces
biomechanical signals to the cells to proliferate and migrate
along its surface. Platelet-enriched plasma is sometimes
added to enhance tissue growth by augmenting the growth
factor profile.12
Similar phenomena are observed with “distraction osteo­
genesis”, where long bones are osteotomized within their
periosteal sleeve and slowly distracted (Fig. 16.3). The
expanding space spontaneously fills with new bone. Just as
in GBR, the environment within the newly created space is
exclusively conducive to bone formation. Both GBR and dis­
traction osteogenesis are examples of what would now be
called endogenous or intrinsic tissue engineering, where
the body grows its own tissue in response to local signals
of trauma, ischemia, inflammation, and mechanical forces
which promote vascularization followed by specialized tissue
regeneration.
Cancellous bone grafting, first described by the New
Zealand plastic surgeon Rainsford Mowlem,13 is an even
earlier example of a tissue-engineering concept, in this case
using the extrinsic principle where a scaffold loaded with cells
is implanted into a new site. Most of the cells probably die
but the bone matrix survives and acts as a scaffold with the
appropriate signals to attract new blood vessel and bone-
forming cells into its interstices and replace it by “creeping
substitution”. The bone remodels according to the biome­
chanical forces of its new environment, as articulated by
Wolf’s law. Bone regeneration has also been encouraged by
the implantation of synthetic porous scaffold materials such
as Proplast and more recently MEDPOR, usually primed with
autologous blood to attract ingrowth of endogenous tissue.
As stated previously, tissue engineering and regenerative
medicine are conceptually different but both disciplines now
borrow from each other’s progress.

233
Components of tissue engineering
induced pluripotent stem cells. Here we will discuss embry­
onic stem cells and then focus on stem cell types most relevant
to the plastic surgeon, the sources of cells, their advantages
and disadvantages and briefly discuss their use in tissue
engineering.
Embryonic stem (ES) cells
Embryonic stem (ES) cells are totipotent, infinitely prolifera­
tive, and can be differentiated into all tissue types. However,
they are also unstable and form teratomas. Additionally, their
sourcing and utilization are limited by ethical/legal concerns.
Successful differentiation protocols have been found to induce
ES cells along specific lineage pathways from all germ layers
towards many specific tissues and organs.14 These cells are
probably immunogenic and ethical issues will persist. There
are many regulatory and organizational issues that will make
pathways into the clinic difficult for elective procedures such
as tissue engineering, especially proving that they will not
form cancer.
Adult stem cells
Adult stem cells are multipotent and limited in their prolifera­
tion capacity and differentiation potential. They are collected
and expanded from tissue biopsies through variations on a
process referred to as the colony forming unit (CFU) assay.15
These progenitor cells are present at low frequencies in the
patient samples, often approximately 1 progenitor cell/100 000
nucleated cells (though the prevalence of adipose-derived
stem cells is significantly higher). The tissue sample is pro­
cessed to remove some of the contaminating cells, and the
remaining mononuclear cells are then plated onto tissue culture
plastic. The progenitor cells are then isolated and purified by
their ability to adhere to the substrate and proliferate. Non-
adherent cells are removed and the medium is supplemented
with factors that promote the survival and growth of the cells
of interest. Over two to three passages the experimenter is left
with a relatively pure population of stem cells.
Two of the first documented adult stem cells were both
found in bone marrow: (1) hematopoietic stem cells (HSCs),
which differentiated into the white blood cell population and
(2) mesenchymal stem cells (MSCs), the progenitors of bone,
cartilage, fat, and muscle.16,17 Additionally, endothelial pro­
genitor cells (EPCs) have been isolated and cultured from
adult peripheral blood.18 Furthermore, MSCs and EPCs are
now known to be present not only in bone marrow but also
in fat tissue associated with the microvasculature,19 where
they are known as adipose-derived stem cells (ASCs). In more
recent years, tissue-specific stem cells have now been isolated
from almost every organ as well as mesenchyme and include
cardiac, liver, lung, kidney, brain, breast and others. Here we
focus on the adult stem cells that are of most use to the plastic
surgeon: the mesenchymal stem cell, the adipose-derived
stem cell and the endothelial progenitor cell.
Mesenchymal stem cells
Classically, MSCs were derived from bone marrow and used
as connective tissue progenitors (cartilage, bone, fat, and
muscle). MSCs have the advantage that they do not express
MHC class II markers and several studies have suggested
they are relatively immune-privileged and may be used as
in the context of tissue engineering as cell maintenance and
behavior including growth and regeneration are influenced
by biochemical and biomechanical interplay. Furthermore, the
development of a functional tissue must be vascularized to
ensure survival of the neotissue. Each of these components is
the purview of the tissue engineer.
Cell sources for tissue engineering
Cells used for tissue engineering may be autologous, heterolo­
gous, or xenogeneic and each type may be mature differenti­
ated or in precursor stem cell form. Ideally, cells for tissue
engineering should be our own cells, and endogenously mobi­
lized to participate in the regenerative process, thereby avoid­
ing issues of rejection and ex vivo cell processing or manipulation.
This is in essence wound healing, where key cells, both local
and distant, are mobilized in response to biochemical gradients
and biomechanical forces. Although this endogenous repair is
ideal it is, unfortunately, not often realized.
Instead, early tissue engineering strategies relied on the
collection of the cell type of interest from the patient, ex vivo
cell expansion, and then reimplantation into the patient,
usually in conjunction with a tissue engineering scaffold
support. Many cell types such as chondrocytes for cartilage,
osteocytes for bone, Schwann cells for nerves, and fibroblasts
for ligament and tendon engineering are highly proliferative
in culture and were early tissue engineering candidates. All
these cells have significant proliferative potential in vitro but,
once implanted, assuming they survive, are not expected to
expand further. However, other differentiated cells such as
adult cardiomyocytes, hepatocytes, and adipocytes are infa­
mously difficult to culture and expand in vitro, significantly
limiting the application of this tissue engineering approach.
A second challenge with this early model of tissue engineering
is that collection of the patient’s primary cells requires biopsy
of the patient’s tissue which is, at best, uncomfortable and, at
worst, impossible due to the diseased state of the tissue.
To circumvent these roadblocks, today’s tissue engineers
will more likely utilize stem cells that can be expanded and
differentiated ex vivo. There are multiple types of stem cells
that have been explored including embryonic, adult, and
Fig. 16.2  Spontaneous regeneration of ruptured extensor digiti minimi tendon at
the wrist (lower tendon held with forceps).

234
CHAPTER 16 • Tissue engineering
cultured, grows more rapidly and can be cultured for longer
periods than bone marrow stem cells before senescence. While
bone marrow-derived mesenchymal stem cells are present at
extremely low frequencies, one gram of adipose tissue can
yield 5000 stem cells making fat a much richer source of stem
cells than bone marrow.24 Additionally, some studies have
illustrated that the ASC population may also have low immu­
nogenicity.25,26 Interestingly, the ASCs are not thought to be a
pure cell population, and for tissue-engineering and fat injec­
tion purposes, it is clear that obsessive attempts to isolate the
“pure ASC” may be counterproductive as the other cells,
especially the EPCs, may also contribute to the apparent
functional effects of stem cells.
As with MSCs, cellular therapies utilizing ASCs have
shown promising results. Partially purified ASCs which still
retain CD34 and CD90 markers grown into spheroid clusters
in a gel culture system differentiate into endothelial cells and
capillary structures in methyl cellulose and these produce
high levels of VEGF. They also rapidly differentiate into adi­
pocytes in adipogenic media, making them potentially suit­
able for adipose tissue engineering.27 Other growth factors
released from ASCs include macrophage colony-stimulating
factor, GMCSF, VEGF, anti-apoptotic factors, hepatocyte
growth factor, and transforming growth factor-β (TGF-β1). In
hypoxia, the VEGF levels are significantly increased, leading
to angiogenesis and decreased apoptosis.
Endothelial progenitor cells
A major requirement in tissue engineering is the incorpora­
tion of a functional vasculature network in the neotissue. One
cell type that shows promise in this arena is the endothelial
progenitor cell (EPC). These cells were first identified in 1997
by Asahara et al.,28 are present in adult circulation, and autolo­
gous cells can be isolated and expanded from peripheral
blood collected through simple venipuncture. Subsequent
research has illustrated there are actually two distinct EPC
populations that participate in vascular repair and angiogen­
esis via different mechanisms. The first population, referred
to as circulating angiogenic cells (or colony forming unit–Hill
cells) are not believed to directly incorporate into the endo­
thelium; instead, they are posited to support resident endo­
thelial cells through paracrine signalling. Endothelial colony
forming cells (ECFCs), on the other hand, have been shown
to differentiate into cells with endothelial phenotype and to
regenerate an endothelial cell population.18 Both cell types can
be used to promote autologous endothelialization.
Challenges associated with adult stem cells
Adult stem cells are advantageous in tissue engineering as
they provide an autologous and/or non-immunogenic source
of cells. However, they are not without limitations. Adult stem
cells display patient-to-patient variations in their prevalence,
proliferative capacity, and differentiation potential.29–31 Addi­
tionally, their utility is also a factor of age and disease state of
the donor. Furthermore, during ex vivo expansion, cells may
exit the cell proliferation cycle (prematurely senesce) or pre­
maturely lose differentiation potential.32,33
Induced pluripotent stem cells
One of the most exciting discoveries in stem cell science
occurred in 2006 when Takahashi et al.34 described a method
allografts. This is a potentially attractive concept where cells
could be sourced from stem cell banks for tissue engineer­
ing.14,20 Due to these attractive properties, MSCs have become
a favored source of cells for the engineering of mesenchymal
tissues, the main building blocks employed for plastic and
orthopedic reconstruction.
Due to their immune-privileged status, many researchers
are attempting to expand upon the utility of MSCs. One
avenue of research is differentiating these cells into cell types
that originate from different germ layers. Furthermore,
researchers have searched for MSCs in other more readily
accessed tissues that can be isolated through less invasive
means. Mesenchymal stem cell-like cells have been observed
in a variety of tissues including skin, muscle, umbilical cord,
placenta, corneal stroma, and dental pulp of deciduous teeth.
However, all of these sources produce progenitor populations
with nuanced differences in their proliferation and differentia­
tion potential with bone marrow-derived MSCs remaining the
most understood of MSC-like cells available.
MSCs injected as a form of cellular therapy have been
widely explored for a variety of conditions including cardiac
disease, peripheral vascular disease, and irradiation injury in
humans. Additionally, a large amount of experience in the
horseracing industry boasts survival, differentiation, and
function of such cells. However, few of these studies were
controlled and tracking of cells was absent or difficult to
assess. With the exception of brain, almost no cases could
confirm that the MSCs survived in significant numbers or that
these survivors differentiated into the injured tissue type.
Interestingly, there is a growing body of knowledge that sug­
gests MSCs injected as a form of therapy do not aid recovery
through direct incorporation with damaged tissue. Most
studies now concede that the paracrine-growth factor-
hormonal-cytokine immune-modulatory effects probably
account for the benefits seen with these stem cells. For
example, ischemia increases homing of these cells to the
injured site where MSCs release high levels of vascular endo­
thelial growth factor (VEGF) to modulate the repair of capil­
laries. Additionally, MSCs injected intravenously in cardiac
infarct models do not implant in the heart nor become heart
tissue but lodge in the lung, where they are activated to secrete
the anti-inflammatory protein TSG-6,21 and it is probably the
anti-inflammatory factor that induces the beneficial effects.
More recently, MSCs have been used as intravenous infusions
for the treatment of multiple sclerosis, with promising results.
It is felt that the response was anti-inflammatory in nature due
to the release of high numbers of M2 anti-inflammatory
macrophages.22 Despite the clear evidence that these cells can
be used in beneficial cellular therapies, the exact mechanism
of action and the best method of employing the cells are still
unclear. However, the clinical use of these cells will continue
to increase. In fact, the National Institutes of Health of the
United States has created a bone marrow stromal cell trans­
plantation program prepared from bone marrow of healthy
subjects to treat patients with inflammatory bowel disease,
cardiovascular disease, and acute graft-vs-host disease.23
Adipose-derived stem cells
Due to their abundance and ease of harvest by liposuction,
adipose-derived stem cells are a preferred autologous stem
cell source. This cell type has similar properties to bone
marrow-derived stem cells but is more abundant, more easily

235
Components of tissue engineering
inflammatory cytokines and local angiogenic growth factors
stimulate endothelial cell migration, proliferation, and mito­
genesis and growth factor cytokines signal distant precursor
cells, including EPCs.10,38 This knowledge has direct impact on
the field of tissue engineering, as it is necessary to know the
appropriate soluble signaling molecules to maintain cell via­
bility both during culture and in vivo, to maintain cellular
phenotype, and to drive lineage specific differentiation of
stem cells. However, biomolecules such as growth factors are
expensive, have short half-lives and angiogenic growth factors
have the particular risk of reactivating dormant cancer when
used for post-cancer resection reconstruction. The need for
local delivery of specific biomolecules at specific concentra­
tions is key to successful tissue regeneration and will be
discussed below via biomaterials with controlled release
properties.
Matrix signals
In addition to surface receptors that engage with soluble
factors, cells also possess receptors such as integrins and
syndecans that bind to a variety of ligands present in the
extracellular matrix (ECM). Many cell types are adhesion
dependent, meaning that if they do not experience biologi­
cally specific binding to an external matrix they will not
maintain viability. Additionally, the number and strength of
bonds with the external matrix affects a wide variety of cel­
lular behaviors ranging from adhesion, focal adhesion forma­
tion and migration to morphology. Additionally, the ECM
binds and sequesters growth factors that also drive cellular
function. Knowledge of the biologically specific binding that
occurs between cells and their matrix has resulted in the
development of biofunctionalized materials that can promote
specific cell functions.
Intercellular signals
Cells also interact with neighboring cells in both native tissue
and in ex vivo culture. These interactions can occur through
two main methods: direct contact via receptors such as cad­
herins and soluble signaling through paracrine factors. In
traditional cell culture, single populations of cells are grown
in isolation, implanted and rely on recruitment of supporting
cellular structures and matrix to evolve into a stable, func­
tional tissue. However, co-culturing cells with other types
prior to implantation can facilitate their survival and function,
such as endothelial cells for vasculature,39 MSCs for paracrine
effects or for their ability to incorporate into developing
tissues in vivo like blood vessels19 and beating cardiomyocytes
with ASCs in vitro to initiate differentiation into cardiac
lineage.40
Mechanics and structure of the environment
In addition to biological signaling that arises directly from a
receptor binding to a ligand, cells also respond to the mechani­
cal properties and dimensionality of their environment. The
cell measures the flexibility of its substrate cytoskeleton
myosin push/pull on the focal adhesion. The response then
leads to changes in tyrosine phosphorylation to signal trans­
duction to the nucleus. In vivo, this phenomenon is commonly
observed during events such as embryonic development and
would healing41 as synthesis/degradation of ECM proteins,
of converting murine somatic fully differentiated skin cells,
mostly fibroblasts, to an embryonic-like state using four
transcription factors. The achievement was then repeated in
2007 for human cells. They are termed induced pluripotent
stem (iPS) cells and possess unlimited proliferation capacity
and the ability to differentiate into cells from all germ layers
both in vitro and in vivo. Additionally, these cells can be
patient-specific, thus present no issues of allogenicity and
they are not encumbered by the same legal/ethical concerns
as embryonic stem cells. In 2012, human iPS cells were even
produced from exfoliated epithelial cells present in urine.35
The major problem associated with the production of iPS cells
is that it requires genetic manipulation of the cells. Further­
more, two of the genes used in this process (c-Myc and KLF4)
are oncogenes.
Much of the research related to iPS cells in recent years has
revolved around improving safety Specifically, inducing a
pluripotent state without genetic modification has been a
major goal. Zhou et al. accomplished this task in 2009 by
delivering the four proteins that the above-mentioned genes
code for directly into the cell. Generation of these cells, referred
to as protein-induced pluripotent stem cells (piPSCs), bypasses
the need for viral or plasmid transfection and reduces the
risks of cancer formation.36 A current drawback to this tech­
nique is that the efficiency of the protein induction is very low,
and research is addressing this limitation. Mice have been
cloned from safer versions of their own iPS cells by implanta­
tion into an enucleated ovum, and these animals have
reproduced and no cancers have been detected.37 Another
technological challenge for the utilization of iPS cell technol­
ogy is that these cells are not immune privileged requiring
collection, induction, expansion, and differentiation of autolo­
gous cells. This is a costly and time-consuming scenario. The
alternative to using autologous cells is to create a bank of iPS
cells that can be HLA matched to the patient. Clinical utiliza­
tion of these cells is being explored. Beyond the realm of tissue
engineering, models based on these cells have also become
powerful tools for drug discovery and investigations into
genetically based diseases.
Cellular interactions with their environment
Many cell types are exquisitely sensitive to stimuli present in
the environment. These stimuli are diverse and include
soluble molecules, molecular recognition sites present in the
solid phase (ECM or biomaterial), interactions with other
cells, substrate stiffness and the micro/nano-structure of the
surroundings. In this section we consider how these param­
eters can affect the phenotype and behavior of cells and in the
biomaterials section we discuss how this knowledge can be
translated into next generation materials for use in tissue
engineering.
Soluble signals
Many soluble biomolecules such as metalloproteins, growth
factors, and chemokines play vital roles locally and systemi­
cally in repair and tissue development. Starting from the
earliest stages of life, embryogenesis is critically dependent
on growth factor cascades released sequentially according to
a genetic code and time clock. In adults these cellular signal­
ing molecules continue to play crucial roles. For example,

236
CHAPTER 16 • Tissue engineering
Biomaterials used in tissue engineering
In addition to cells, a main component of solid tissues is the
extracellular matrix which provides structure and biochemi­
cal signals to the cells. However, when cells are expanded
outside the body, they typically grow in monolayers, not the
intricate patterns of a fully realized three-dimensional tissue.
Therefore, cells are often seeded onto a 3D scaffold. In this
light, biomaterial scaffolds can be thought of as artificial
ECMs. Here we discuss the main classes of biomaterials for
tissue engineering and methods employed to create materials
that drive desired cellular behaviors.
Many criteria exist for biomaterials to be suitable for tissue
engineering applications, as outlined in Box 16.1. A wide
movement of cells, or fluid shear from blood flow modulate
the mechanical properties of the cells’ environment.
The mechanical properties of tissues and their components
vary widely (Table 16.1 and Fig. 16.4A) and so the physical
microenvironment of particular cells in vivo also varies signifi­
cantly. In traditional cell culture, cells are grown on rigid and
two-dimensional tissue culture plastic, the equivalent of cells
being firmly stretched in clinical situations. They then move,
migrate, and proliferate until they reach confluence. The same
cells when placed in a soft 3D culture gel environment sense
a modulus that encourages the cells to cluster together into
spheroids (Fig. 16.5). When the cells pull on the soft matrix
they pull it into a ball, stop dividing, and many apoptose. If
they are stem cells the stiffness of the environment may
promote lineage-specific differentiation.42–44 Engler et al.44,45
have shown that MSC lineage specification can be altered in
vitro when cells are cultured on substrates of different stiffness
(see Fig. 16.4). Cellular morphology, gene expression, migra­
tion, and proliferation are also modulated by the mechanical
properties of their microenvironment.43,46–48
The response of cells to environmental stiffness and dimen­
sionality provides the tissue engineer with an additional tool
that can be used to direct cell function. For example, chondro­
cytes change morphology and lose their chondrogenic capac­
ity in 2D monolayer culture but can maintain these features
in 3D culture.49 The response of cells to mechanical stimula­
tion is used in tissue engineering strategies such as distraction
osteogenesis. Periods of stretch promote proliferation and
migration while relaxation incites cells to cluster together and
terminally differentiate into bone. Guided tissue regeneration
similarly works by presenting the cells with a hard surface
that stimulates them to divide and to migrate.
Stem cell mechanobiology is attracting increased attention
as understanding develops of the influence of substrate
mechanics and architecture on differentiation.45 Cells may be
cultured in 3D on biomaterial supports such as porous scaf­
folds, hydrogels, or microspheres under conditions designed
for the desired cell attachment, migration, proliferation, and
differentiation.50,51 This knowledge can influence the develop­
ment of new biomaterials with tailored properties and direct
the fabrication of these materials into three-dimensional scaf­
folds to maximize the healing process.
Table 16.1  Typical mechanical properties of selected tissues
and tissue components
Material
Elastic
modulus
Yield
stress
Maximum
strain
Elastin
100 kPa
300 kPa
300%
Cartilage
10 MPa
8–20 MPa
70–200%
Skin
35 MPa
15 MPa
100%
Muscle
fascia
350 MPa
15 MPa
170%
Tendon
700 MPa
60 MPa
10%
Cortical
bone
17.4 GPa
160 MPa
1.8%
(Adapted from Palsson BO, Bhatia SN. Tissue Engineering. Upper Saddle River,
NJ: Pearson Education; 2004; and Cowin SC. The mechanical properties of
cortical bone tissue. In: Cowin SC (ed). Bone Mechanics. Boca Raton, FL: CRC
Press; 1989.)
Fig. 16.4  Effects of substrate mechanical properties on differentiation. (A) Range
of microenvironment elasticity for various solid tissues. (B) Mesenchymal stem
cells (MSCs) grown on collagen I-coated hydrogels of various elastic moduli
differentiate to phenotypes corresponding to tissues of the same elasticity as the
hydrogels. (Adapted From Engler AJ, Sen S, Sweeney HL, et al. Matrix elasticity
directs stem cell lineage specification. Cell. 2006;126:677–689.)
A
B
Elastic substrate (Ε)
8–17 kPa
25–40 kPa
0.1–1 kPa
96 hrs
24 hrs
4 hrs
h
Collagen I
10% serum
Solid tissues
Blood
Fluid
Brain
Muscle
Collagenous bone
1 kPa
10 kPa
100 kPa = Ε,
Elasticity of
micro-
environment
MSC

237
Components of tissue engineering
clinical acceptance, with very high costs and relatively low
rates of translation from the laboratory to the clinic.53 Despite
this, much research continues and a range of such materials
are in this testing pipeline.
Biomaterials for tissue engineering may be naturally occur­
ring in the human body or derived from other natural or
synthetic sources. They may be ceramics, polymers, hydro­
gels, or composites of these. Decellularized tissues are also
used as biomaterials to support tissue growth. The physical
form of biomaterials for tissue-engineering applications can
also vary to suit the application – solid materials, porous
scaffolds, microspheres, hydrogels, injectable materials that
may cross-link in situ, or other configurations can be fabri­
cated. Whilst it may be feasible to produce complex biomimetic
materials to influence cell fate ex vivo, simplicity is also desir­
able to facilitate regulatory approval and translation of new
materials into clinical application.54 Selection of biomaterials
will depend on the specific requirements of the tissue being
targeted.
Biodegradable materials
Most biomaterials used in tissue engineering applications are
biodegradable so that, over time, the material will disappear
at a specified rate and be replaced by neotissue. The rate of
degradation can vary widely so the mechanical integrity may
be retained for as little as days or weeks, or as long as a year
or more. The rate of degradation and loss of integrity will
depend on the type of biomaterial, the site of implantation,
properties of the biomaterial construct such as surface area to
volume ratio, size, and surface chemistry. Fig. 16.6 shows
examples of the loss of mechanical strength of biodegradable
polymers.
As these materials degrade over time, one of the major
challenges with biomaterial design is the prevention of sudden
loss of physical integrity. Additionally, rapid degradation may
lead to excessive concentrations of the degradation products
of a biomaterial that can cause adverse tissue reactions. This
variety of synthetic and naturally derived biomaterials have
been considered to meet these needs although they do not
generally ideally satisfy all of these criteria.52 Early tissue-
engineering attempts tended to utilize materials that had
already been used in other clinical applications, such as the
polymers used in degradable sutures, or natural materials
such as coral, alginate and collagen. More recently, numerous
alternatives have been formulated as researchers aim to
develop multifunctional biomimetic biomaterials to suit the
specific demands of tissue engineering, which can differ sig­
nificantly from those of other biomaterial applications. New
materials require extensive testing in vitro and in vivo before
Fig. 16.5  (A) Preadipocytes growing on two-dimensional tissue culture plastic.
(B) The same cells growing in three-dimensional gel culture form spheroids.
(Reproduced from Stillaert FB, Abberton KM, Keramidaris E, et al. Intrinsics and
dynamics of fat grafts: an in vitro study. Plast Reconstr Surg. 2010;126:
1155–1162.)
A
B
BOX 16.1  Criteria for biomaterials to be suitable for tissue
engineering applications
Biocompatibility, i.e., eliciting a suitable host response in the specific
application. Suitable mechanical properties for target tissue and
implantation site.
Biodegradability over desired timeframe with suitable mechanical
properties and non-harmful degradation products.
Not containing harmful impurities.
A host response in vivo that does not prevent desired tissue
development; i.e., inflammation and foreign-body response are
not excessive.
Ability to be fabricated into desired structures.
Cost-effective.
Ability to be sterilized without harmful changes to their chemical or
physical properties.
Adequate stability and shelf-life.
Ability to promote desired cellular responses, e.g., proliferation,
differentiation, gene expression.

238
CHAPTER 16 • Tissue engineering
those that have not yet been identified. One interesting appli­
cation of dECM is in the expansion of stem cells. Recently,
MSCs have been cultured on dECM and the cells have illus­
trated significantly higher proliferation rates and maintenance
of stem cell phenotype, clearly illustrating the importance of
environmental control during cell culture conditions.57
Polymeric biomaterials
Hydrophobic polymers
Numerous synthetic polymers are used in medicine and some
of these are biodegradable (or bioresorbable) in vivo. Fig. 16.7
shows examples of the chemical structures of biodegradable
polymers that could be used in tissue engineering. Among
these, the polyesters poly(glycolic acid) (PGA), poly(lactic
acid) (PLA) and poly(ε-caprolactone) (PCL) and their copoly­
mers such as poly(lactide-co-glycolide) (PLGA) have been
widely used in tissue-engineering approaches.58 The mechani­
cal strength and degradation rate of these polymers can be
tuned over a wide range by changing the polymer properties
(molecular weight, composition, molecular architecture,
crystallinity, hydrophobicity).59 However, many biodegrad­
able synthetic polymers are quite hydrophobic, unlike the
ECM. Exposure of such surfaces in vivo can lead to nonspecific
adsorption of proteins of mixed orientation and states of
denaturation, potentially resulting in a foreign-body reaction
(FBR) to the material.60
Hydrogels
Hydrogels are water-swollen cross-linked polymer networks
which can absorb up to thousands of times their dry weight
of water, which gives them the advantage of being more like
most natural tissues and allowing mass transport to and from
cells.61 They may be formed by physical (hydrogen bonds,
ionic interactions, molecular entanglements, hydrophobic
interactions) or chemical interactions (covalent bonds) and
this affects their stability and degradation rates. The mechani­
cal strength of hydrogels can be tuned within a limited
range and their degradation tends to be faster than for other
biomaterials. Hydrogels may be naturally derived (e.g.,
collagen, gelatin, hyaluronic acid, alginate) or synthetic (e.g.,
poly(ethylene glycol) (PEG)-based polymers).
Ceramic biomaterials
Ceramic biomaterials are primarily utilized in tissue engineer­
ing of hard tissues. Calcium phosphates, such as hydroxyapa­
tite, and bioactive glasses have been developed as bioceramics
for bone tissue engineering. Bioceramics are brittle but have
high compressive strength, can bond strongly to bone, and
can be osteoinductive.
Advanced biomaterials for tissue engineering
In the early days of tissue engineering, commodity materials
were often used. However, these materials are often not suit­
able, as described above. Now we are in an exciting age of
biomaterials science where materials are being designed to
have tailored interactions with biology using our growing
knowledge about how cells interact with their environments
in order to lead to improved healing.62 Here we present an
overview of some of these strategies and how they apply to
tissue engineering.
is well known for the polylactides and polyglycolides that can
undergo autocatalytic degradation with decreasing pH inside
a thick-walled construct and “acid burst” leading to reduced
local pH and tissue necrosis. For this reason, it is important
to consider degradation rates and mechanisms in designing
the composition and architecture of tissue-engineering con­
structs using biomaterials.
Natural biomaterials
Natural biomaterials may have the benefits of being chemi­
cally similar or identical to molecules in the body, readily
degraded in vivo, and able to interact with cells on a molecular
level. However, they can be difficult to obtain and purify, vary
in properties between batches, may be difficult to sterilize,
alter their properties during storage, and they may elicit sig­
nificant immunogenic responses. Naturally derived biomate­
rials investigated for tissue engineering include proteins (e.g.,
collagen, gelatin, silk), polysaccharides (e.g., chitosan, hyal­
uronic acid), polynucleotides and extracts of ECM compo­
nents.55 These materials may be modified to tailor their
properties, such as via chemical cross-linking to slow degra­
dation and increase mechanical strength.
One natural biomaterial that is gaining increasing attention
is decellularized extracellular matrix (dECM). During decel­
lularization, cells are removed from allografts or xenografts
to reduce immunogenicity but much of the complex composi­
tion and architecture of the ECM may be retained.56 One limi­
tation of synthetic materials is that they cannot contain the
large number of biomolecules present in the ECM whilst the
dECM has the potential for containing important factors, even
Fig. 16.6  Changes in tensile strength of biodegradable polymers during
degradation in vitro. (Reproduced from Suzuki M, Ikada Y. Biodegradable polymers
in medicine. In: Reis RL, Román JS (eds). Biodegradable Systems in Tissue
Engineering and Regenerative Medicine. Boca Raton: CRC Press, 2005.)
Time in phosphate buffered saline at 37ºC in vitro (weeks)
PGA (polyglycolide)
MONOCRYL (Glycolide-ε-caprolactone copolymer)
BIOSYN (glycolide-p-dioxanone-trimethylenecarbonate copolymer)
MAXON (glycolide-trimethylenecarbonate copolymer)
P(LA/CL) (L-lactide-ε-caprolactone copolymer) 75/25
PDS (poly-p-dioxanone)
PLLA (poly-L-lactide)
0
10
20
30
40
50
60
70
80
90
100
Strength remaining (%)
1
0
2
3
4
5
6
7
8
9
10
11
12

239
Components of tissue engineering
biomaterial constructs such as within the walls or on the
surfaces of scaffolds or hydrogels.47 It is critical to the design
of an effective delivery system that the payload of bioactive
molecules retains its conformation and bioactivity throughout
the loading and release processes. An additional challenge is
the prediction of release rates in vivo, which generally differ
significantly from those in vitro.
Controlled delivery of growth factors is desirable as they
have short half-lives in vivo, are relatively slow to diffuse due
to their size, and may have undesired effects if delivered
systemically.46 The potential benefits of growth factor delivery
on adipose tissue engineering were demonstrated in vivo
using gelatin microspheres to deliver fibroblast growth
factor-2 (FGF-2) over an extended time in a mouse tissue-
engineering chamber.68 Chambers containing an autologous
Tailored delivery systems
As discussed above, soluble bioactive molecules play a critical
role in natural tissue morphogenesis and control of their
delivery both spatially and temporally to growing tissue may
significantly enhance tissue-engineering outcomes. Molecules
such as growth factors, anti-inflammatory peptides, and
drugs may be incorporated into biomaterial delivery vehicles
for release at the desired time during tissue development.63–65
The release systems may be designed to deliver multiple
molecules over different timescales via continuous or pulsatile
delivery,66 which may be programmed or triggered by some
change in the local environment.67 Delivery systems can be
fabricated from biodegradable polymers in the form of micro-
or nanoparticles or capsules, or incorporated into other
Fig. 16.7  Chemical structures of selected biodegradable polymers of potential interest in tissue engineering. (Reproduced from Kohn J, Abramson S, Langer R.
Bioresorbable and bioerodible materials. In: Ratner BD, Hoffman AS, Schoen FJ, et al. (eds). Biomaterials Science. Amsterdam: Elsevier, 2004;115–217. Chitosan
reproduced from Guibal E, Vincent T, Blondet FP. Biopolymers as supports for heterogeneous catalysis: focus on chitosan, a promising aminopolysaccharide. In: SenGupta AK
(ed). Ion Exchange and Solvent Extraction: A Series of Advances. Boca Raton: Taylor & Francis; 151–292. Hyaluronic acid reproduced from Yannas IV. Natural materials. In:
Ratner BD, et al (eds). Biomaterials Science – An Introduction to Materials in Medicine. Amsterdam: Elsevier; 2004.)
Polydioxanone
n
O
O
O
C
(CH2)2
CH2
Poly(β-hydroxybutyrate)
n
O
CH2
CH
CH3
O
C
Poly(hydroxyvalerate)
n
O
CH2
CH
O
C
CH2
CH3
Poly(ε-caprolactone)
n
O
(CH2)5
O
C
Poly(lactic acid)
n
CH
CH3
O
O
C
Poly(glycolic acid)
n
CH2
O
O
C
Poly[(p-methyl phenoxy)
(ethyl glycinato) phosphazene]
O
O
P
HN
N
O
n
Poly(PCPP-SA anhydride)
bis(p-carboxyphenoxy)propane (PCPP)
sebacic acid (SA)
O
O
C
O
O
C
O
O
C
O(CH2)3
(CH2)8
O
C
n
m
Chitosan
O
O
*
4
1
CH2
NH2
OH
x
O
O
4
1
CH2
NH
CH3
C
O
OH
y
Hyaluronic acid
n
O
O
*
O
*
4
1
CH2
NH
OH
OH
O
O
O
4
1
OH
OH
OH
O

240
CHAPTER 16 • Tissue engineering
features that lead to reduced fibrosis (e.g., porous polymers
with pores of around 30 µm74 or small-diameter fibers75).
Surface chemical properties can also be used to reduce the
FBR to biomaterials, such as via attachment of oriented
biomolecules.76
One of the more successful strategies is the use of chemical
surface modifications to create the so-called non-fouling sur­
faces. An initial stage of FBR is the adsorption of a complex
layer of biomolecules from body fluids that can be denatured
and lead to an immune reaction. Non-fouling materials (or
stealth materials) resist the adsorption of these proteins.
Usually these surface modifications make the material hydro­
philic yet electrically neutral and the mechanism of action is
believed to be that the surfaces hold their waters of hydration
instead of allowing them to be displaced by adsorbing bio­
macromolecules. A myriad of strategies have been employed
in order to generate non-fouling outer layers on biomaterials
(e.g., by layer-by-layer assembly of polyelectrolytes ending
with a nonadhesive outer layer,77 or attachment of PEG15,78).
However, many strategies that show promise in vitro suffer
from enzymatic attack and changes to the biomaterial surface
in vivo during the acute inflammatory stages of the wound-
healing response, and so their in vivo performance may
be disappointing. Creating new generations of non-fouling
materials with improved in vivo performance is an active area
of research and includes candidates such as zwitterionic
polymers, mixed charged polymers, and polyoxazolines.79
Fig. 16.8  Effect of controlled release of a growth factor on tissue development in
vivo. Cross-linked gelatin microspheres suspended in collagen were used to deliver
fibroblast growth factor-2 (FGF-2) in a tubular mouse tissue-engineering chamber
enclosing superficial epigastric vessels in the groin of mice, with (+) or without (−)
an autologous fat graft. (From Vashi AV, Abberton KM, Thomas GP, et al. Adipose
tissue engineering based on the controlled release of fibroblast growth factor-
2(FGF-2) in a collagen matrix. Tissue Eng. 2006;12:3035–3043.)
0.000
Collagen +
free FGF-2 (-)
Collagen +
FGF-2-loaded
microspheres (-)
Collagen +
FGF-2-loaded
microspheres (+)
Collagen +
blank microspheres(-)
Collagen +
blank microspheres(+)
Collagen +
free FGF-2 (+)
0.001
0.002
0.003
0.004
0.005
0.006
0.007
0.008
Volume displacement (g)
Fig. 16.9  Potential reversible responses of “smart polymers” to environmental
stimuli. (A) Gelation. (B) Surface adsorption. (C) Collapse on to surface.
(D) Swelling or contraction of hydrogel. (From Hoffman AS. Applications of “smart
polymers” as biomaterials. In: Ratner BD, Hoffman AS, Schoen FJ, et al (eds).
Biomaterials Science. Amsterdam: Elsevier, 2004;107–115.)
A
B
C
D
+ Stimulus
Gelation
Surface
adsorption
Collapse
onto surface
Swelling or
contraction
of hydrogel
– Stimulus
+ Stimulus
– Stimulus
+ Stimulus
– Stimulus
+ Stimulus
– Stimulus
fat graft and gelatin microspheres loaded with FGF-2 grew
significantly more tissue than controls with free FGF-2 or
blank microspheres (Fig. 16.8).
Similar approaches are also being investigated for cell
delivery, wherein cells may be loaded onto degradable micro­
spheres in vitro, allowed to attach, migrate, and proliferate on
or within the spheres and then delivered to a target site for
tissue growth. This significantly increases the surface area
available for attachment of adherent cells and using biode­
gradable beads that can be delivered in vivo avoids the need
to detach the cells from the culture surface for delivery.69 It
has been shown that ASCs loaded into chitosan microspheres
were viable and maintained their multipotency after in vitro
culture.70
Smart polymers
Changes in environmental conditions can result in changes
to the molecular conformation of many materials. These
environmentally-induced changes may be harnessed, result­
ing in so-called smart polymers (Fig. 16.9). These types of
change can be used to encapsulate and release payloads of
cells or drugs, to form gels upon injection in vivo, or for cell
sheet engineering.60,71,72 A classic example of a smart polymer
is the thermo-responsive polymer N-isopropylacrylamide
(NIPAM) which is often used in tissue engineering applica­
tions. This material can be used to grow confluent cell sheets
and then to detach the intact sheet along with the ECM that
the cells have deposited.73
Non-fouling materials
Another key challenge related to biomaterials is the foreign
body response (FBR) seen in vivo with almost all synthetic
biomaterials used clinically, be they polymers, ceramics, or
hydrogels. Approaches to reduce the FBR to biomaterials in
vivo may use physical properties, such as by creating spatial

241
Components of tissue engineering
involves embedding cells within a hydrogel or incorporating
microporous structures to enable cellular infiltration and
metabolite diffusion. The design of biomaterial constructs for
tissue engineering is complex and varies significantly depend­
ing on the tissue target. Even for specific tissue types, various
design criteria may be considered (Box 16.2) and the defini­
tion of ideal designs is not yet well established.
Biomaterial constructs for tissue engineering can be fabri­
cated into structures such as porous scaffolds and hydrogels,
meshes or microspheres using a variety of techniques, including
polymer phase separation, particle or foam templating, cryo­
gelation, electrospinning and rapid prototyping methods like
3D-printing (discussed further in the emerging technologies
Biofunctionalized materials
Cells will often attach to biomaterials implanted into the body,
and this attachment is mediated by a layer of adsorbed pro­
teins that contain cell binding sites. The abovementioned
non-fouling technologies result in interfaces that do not
adsorb proteins so are also resistant to cellular attachment.
Such interfaces represent a “blank slate” for tissue engineers.
These surfaces can then be decorated with bioactive mol­
ecules, usually through covalent immobilization. These bio­
functionalized materials are then able to specifically interact
with receptors on the cell surface and drive cellular behavior
with biological specificity. The most common biofunctional­
ization strategy is to generate materials that contain ligands
that engage specific integrin receptors. Thus, only cells that
express the appropriate integrin are able to adhere to the
material. The interaction of cells with such a surface depend­
ing on whether the surface has binding sites of suitable ori­
entation and conformation to enable attachment of the cells,
and the density of such sites. This in turn can affect cell
morphology and migration rates (Fig. 16.10). Mostly these
ligands are adhesion sites found in the native ECM (e.g., the
RGD tri-peptide found in fibronectin),80 but researchers have
also performed analyses to discover unique receptors present
on their cell type of choice. For example, a phage display
library was screened in order to identify ligands specific to an
outgrowth of endothelial progenitor cells but not other com­
monly occurring cell types. These ligands were then used to
functionalize a biomaterial resulting in an interface inhabited
exclusively by the target cell of interest.81
In addition to adhesion ligands, other bioactive molecules
have been incorporated into synthetic materials in order to
tailor biological interactions. For instance, the ECM acts to
locally sequester growth factors, and similarly synthetic ECMs
with tethered growth factors have been generated. Addition­
ally, biofunctionalization has been used to tailor biomaterial
degradation. The abovementioned biodegradable materials
usually undergo some form of degradation reaction, for
example hydrolysis of a chemical bond. This does result in
a material that degrades; however, the degradation occurs
regardless of cellular/tissue infiltration. However, cells move
through the native matrix through secretion of matrix metallo­
proteinases that locally degrade the matrix. Advanced hydro­
gels that are crosslinked with peptide sites that are degraded
by such metalloproteinases have been developed so that the
biomaterial degradation rate adapts to the tissue growth rate.48
Tissue engineering constructs
The abovementioned biomaterials are generally fabricated
into a tissue scaffold to support regeneration. Usually, this
Fig. 16.10  Effects of cell–substrate interactions on cell migration
speed. (Reproduced from Palsson BO, Bhatia SN. Tissue
Engineering. Upper Saddle River, NJ: Pearson Education; 2004.)
Intracellular contraction forces
Substratum adhesive strength
Large
Intermediate
Small
Negligible
Significant
Negligible
Typical cell morphology
Migration speed
BOX 16.2  Key design criteria for tissue engineering scaffolds
Biomaterial selection, including:
chemical composition
molecular weight
polydispersity
crystallinity
presence of any impurities (e.g., catalyst residues)
Internal 3D porous architecture:
pore sizes
pore shapes
pore interconnectivity
volume fraction of pores
wall thicknesses
Surface chemistry
Surface topology
Degradation profile and products
Mechanical properties (initially and during degradation):
elasticity
strength
toughness
Effects on cellular processes
In vivo effects (e.g., immune responses, foreign body reactions)
Ability to be fabricated reproducibly and economically
Ability to be sterilized without unacceptable property changes

242
CHAPTER 16 • Tissue engineering
required to fit cells or blood capillaries. Cao et al. found that
pores of at least 300 µm were needed for tissue survival in
PLGA scaffolds where a significant FBR occurred inside the
scaffolds.91
Ideally the mechanical strength of tissue-engineering con­
structs is matched to that of the target tissue to favor the
desired cellular behavior, as discussed above, and ensure the
construct will not collapse under the forces applied in vivo but
will also not cause stress shielding to the surrounding tissue,
which can lead to changes such as bone resorption. The con­
struct should maintain its strength during tissue development
section).51,82–86 The techniques for tissue-engineering scaffold
fabrication have been widely reviewed.53,87–90 A limited but
growing range of biodegradable scaffolds are available com­
mercially for cell culture and very specific approved clinical
applications, examples of which are discussed in the case
studies section at the end of this chapter.
Scaffolds of various internal architectures can be produced
(Fig. 16.11) with different mechanical and degradation prop­
erties. The porosity must be interconnected and of sufficient
size for molecular transport, cell migration, and vasculariza­
tion. The size required may be larger than the minimum size
Fig. 16.11  Scaffold morphologies produced by thermally induced phase separation (A–C). (D) Particulate leaching. (E) Electrospinning. (F) Rapid prototyping. (A–C,
Reproduced from Cao Y, Croll T, Cooper-White JJ, et al. Production and surface modification of polylactide-based polymeric scaffolds for soft tissue engineering. In: Hollander
AP, Hatton PV (eds). Biopolymer Methods in Tissue Engineering. Totowa, NJ: Humana Press; 2004:87–112. (D–E) Reproduced from Kretlow JD, Mikos AG. From material to
tissue: biomaterial development, scaffold fabrication, and tissue engineering. AIChE J. 2008;54:3048–3067. (F) Reproduced from Shor L, Guceri S, Chang R, et al. Precision
extruding deposition (PED) fabrication of polycaprolactone (PCL) scaffolds for bone tissue engineering. Biofabrication. 2009;1:1–10.)
A
B
C
D
E
F

243
Components of tissue engineering
Vasculogenesis, angiogenesis and inosculation
Understanding how the human vasculature develops,
expands and connects to a grafted tissue is fundamental to
successfully building a capillary network in vitro that is able
to link to host’s capillaries in vivo and ultimately support the
growth, expansion and function of a tissue-engineered con­
struct. Vasculogenesis refers to the development of new blood
vessels from progenitor stem cells.98 In the embryo, mesoderm
cells are stimulated by FGF-2 to form hemangioblasts, a
common precursor for vessel and blood cell formation. After
pooling together into blood islands, hemangioblasts located
at the periphery differentiate first into angioblasts and then
into endothelial cells, which start to coalesce resulting in a
primary vascular plexus.99 Though initially it was assumed
that vasculogenesis was exclusive to the embryological period,
we know now that it also takes place in adult life driven by
bone marrow-derived endothelial progenitor cells.100 Angio­
genesis describes the sprouting of blood vessels from pre-
existing ones. Briefly, in this process endothelial cells respond
to an angiogenic signal and increase vascular permeability
allowing the extravasation of plasma proteins that form a
transient matrix. A tip cell then migrates out of the vessel into
the previously deposited matrix and leads the sprouting fol­
lowed by stalk cells, which form a lumen. Once neighboring
sprouts fuse, the neovessel becomes perfused and is further
stabilized by the recruitment of pericytes and basement
membrane restitution.101 Inosculation describes the process by
which capillaries from a grafted tissue connect to those of the
wound bed where it is applied. The exact mechanism by
which such connection occurs and whether it is dependent
on the characteristics of the grafted tissue are not entirely
clear.102,103 What is clear, though, is that because it takes several
days to ensue, inosculation is one of the limiting factors affect­
ing prompt perfusion of a tissue-engineered construct.104–107
Such a period of time is too long to expect anything but very
thin tissues to survive the transplantation anoxia. Neverthe­
less, some studies have achieved faster inosculation by using
either fibroblasts or FGF-2, giving some interesting insights
on how to accelerate the process and eventually make it clini­
cally feasible.108,109
Elements and strategies of vascularization
in tissue engineering
As stated before, engineering of any 3D tissue or organ is
highly dependent on an adequate blood supply. This can be
achieved by either an extrinsic or an intrinsic approach. The
extrinsic approach implies the direct implantation of a seeded
scaffold that is gradually invaded by the host’s vasculature,
meanwhile relying solely on diffusion for survival. Alterna­
tively, through a process named pre-vascularization, a capil­
lary bed can be fabricated in vitro and then implanted in vivo
where it will become perfused by inosculation rather than by
capillary invasion (Fig. 16.13). In the intrinsic approach the
organism is used as a bioreactor so that capillary sprouting
occurs either before or concomitantly with cell differentiation
and tissue growth.110 But before going into the detail of each
of these approaches, one must first be familiarized with the
different basic elements common to both strategies, which,
similar to tissue engineering itself, are: cells, scaffolds and
growth factors.
until the newly formed tissue can withstand the mechanical
forces at play.92 Scaffold mechanics can be varied by adjusting
the scaffold material, fabrication process, porosity (pore size,
volume and architecture), grain size (for ceramics) and using
composites such as nanoparticles of hydroxyapatite in a
polymer matrix, as well as nanoparticle surface patterning.93,94
The effect of porosity in ceramics is illustrated in Fig. 16.12.
Ceramics are prone to brittle failure and purely ceramic scaf­
folds with the porosity desired for effective vascularization
may not have sufficient strength and toughness for load-
bearing bone tissue engineering, which limits their clinical
application.95
Fabrication processes can incorporate viable cells and/or
bioactive molecules with biomaterials by suitably gentle
processes and scaffold-free construct production methods are
also being developed, such as by using magnetic fields to
arrange cells.96
Vascularization in tissue engineering
The remarkable advances in tissue engineering during recent
years have highlighted the importance of vascularization as
it has become clear that survival, growth and function of an
engineered construct are highly dependent on an adequate
and timely blood supply. Generally, cells do not survive
beyond 150 µm from a capillary,6,97 although certain cells and
tissues vary in their oxygen needs. As stated earlier, whilst the
first attempts of tissue fabrication were focused mainly on
tissues of either low oxygen requirements, such as cartilage
and tendon, or very small thickness, such as skin, able to
survive by diffusion, it is now widely accepted that the fabri­
cation of thicker and more “oxygen-demanding” surrogates
cannot be conceived without first addressing the issue of
vascularization, currently one of the major hurdles in translat­
ing laboratory-based tissue engineering products into human
practice.
Fig. 16.12  Ceramic scaffold mechanical properties as a function of porosity.
(From Boccaccini AR. Ceramics. In: Hench JRJ (ed.) Biomaterials, Artificial Organs
and Tissue Engineering. Cambridge: Woodhead; 2005.)
Porosity
0
10
20
30
40
50
60
70
80
90
Modulus (GPa)
0.1
0
0.2
0.3
0.4
0.5

244
CHAPTER 16 • Tissue engineering
that adipose-derived stem cells have been used both as sup­
portive and endothelial cells,113–115 with one study actually
suggesting that ASCs alone may suffice for the formation of
vascular lumina upon implantation.113 Of note also is the role
of fibroblasts. As mentioned earlier, experiments using high
concentrations of these cells have reported accelerated inos­
culation as early as one day post-implantation.108 A further
cellular element that has been explored with success in pre­
vascularization is the stromal vascular fraction (SVF) from
lipoaspirates. The presence of endothelial progenitors within
the SVF cell population, as well as its remarkable proangio­
genic potential, was reported more than a decade ago.116,117
More recently, other studies have reported on the successful
growth of a functional capillary bed using SVF only.103,118 The
current knowledge on the different components of SVF,
namely mesenchymal stem cells, endothelial cells, endothelial
precursors, vascular smooth muscle cells, pericytes and
angiogenic growth factors make this lipoaspirate-derived
cocktail a very reasonable and attractive source of the cellular
elements needed to build a capillary network.119,120
Scaffolds
The characteristics of scaffolds for tissue engineering have
been described above. For vascularization, several of these
matrices have been tested, all having their inherent pros and
cons. Natural scaffolds include collagen, fibrin, starch, Matri­
gel®, decellularized matrix and silk fibrion. Synthetic materials
on the other hand comprise mainly polyethylene glycol, poly
(lactic glycolic) acid and polyurethane.121 Furthermore, clini­
cally approved dermal substitutes such as Integra® and Mat­
riderm® have also been used for prevascularization purposes
by seeding them with endothelial colony forming cells
(ECFCs) in association with either human dermal fibroblasts
(hDFs) or bone marrow-derived mesenchymal stem cells
(BMSCs).122 Whatever scaffold is used, it is important to
consider the close cell–cell and cell–matrix interactions as well
as the mechanical features that eventually guide and stabilize
the newly formed vessels.123 Likewise, since vasculogenesis
and angiogenesis are deeply influenced by the presence of an
Elements
Cells
A number of different cell sources have been trialed with
varying degrees of success. In general terms, the vast majority
of experiments attempting to build a capillary network use
two cell lines, namely endothelial cells and supportive cells
(Table 16.2). The logic in this is to try to mimic to some extent
the processes of vasculogenesis and angiogenesis in which the
lumen is formed by the endothelial cells while the stabiliza­
tion and maturation of the newly formed vessels is partly the
result of the structural and paracrine action of the perivascular
supportive cells.
As can be seen in Table 16.2, there are a number of possible
combinations. Whatever mixture is chosen, most of the latest
research points towards better results obtained with bi-culture
or even tri-culture compared with seeding endothelial cells
alone.104,108,111,112 Additionally, it is interesting to note the fact
Fig. 16.13  Schematic illustration of the vascularization of a grafted tissue by
inosculation and capillary invasion (angiogenesis). When a vascular network is
fabricated prior to implantation, inosculation between the scaffold’s vessels and bed
capillaries at the periphery of the construct provides blood supply to the entire
block. Conversely, if the scaffold is implanted without a pre-formed vascular
network, adequate blood supply is only provided once host capillaries have
completely invaded the construct.
Inosculation
Angiogenesis
Table 16.2  Cells most commonly used for building capillary
networks.
Endothelial cells (ECs)
Supportive cells
Human umbilical vein ECs (HUVEC)
Human dermal microvascular ECs
(HDMEC)
Human endothelial colony-forming
cells
Human iPS-derived endothelial
precursor cells
Adipose-derived stem cells
Blood-derived endothelial
progenitor cells
Adipose-derived endothelial
progenitor cells
Bovine pulmonary artery ECs
Rat aortic ECs
Smooth muscle cells
Pericytes
Fibroblasts
Bone marrow MSC
Adipose-derived stem
cells
Human embryonic
stem cells – MC
Human iPS –
mesenchymal
precursor cells
10 T1/2 embryonic
murine mesenchymal
stem cells
iPS, induced pluripotent stem cells; MSC, mesenchymal stem cells; MC,
mesenchymal cells.

245
Components of tissue engineering
Strategies
Extrinsic vascularization model of tissue engineering
The model of extrinsic vascularization relies on diffusion
immediately after implantation until vascularization invades
the construct from the periphery through into its core, a
process that can take several days or weeks to be fully estab­
lished. This model was the traditional prototype for 3D tissue
engineering from its inception; nevertheless, it has become
somewhat obsolete due to the growing knowledge on the
importance of an adequate and prompt blood supply. Nowa­
days, it is mostly conceded that reliance on a 3D prefabricated
structure to survive implantation by the random invasion of
capillaries from the implantation bed is unlikely to be a viable
concept, especially for tissues sizeable enough to be clinically
useful. It may, however, still have a role in the engineering of
thin or avascular tissues such as skin or cartilage. A number
of animal models exploiting the extrinsic principle have been
described (Box 16.3).
Interestingly, some studies have suggested that vasculo­
genesis in the subcutaneous model requires a significantly
higher number of cells and growth factors compared with the
brain, a fact that highlights the importance of implanting the
scaffold in a well-vascularized environment to optimize per­
fusion.112 To try to speed up the process, a number of modifica­
tions have been trialed to facilitate rapid vascularization upon
implantation, including angiogenic growth factors,127,135 endo­
thelial cells and MSC co-culturing for their angiogenic and
paracrine effects,136,137 patterning of synthetic or collagen
matrices with artificial vascular channels,138–141 bioactive scaf­
folds containing growth factors for controlled release,142–145 or
decellularized whole-organ tissues that retain a skeletonized
vascular tree which may be seeded with endothelial cells ex
vivo or spontaneously endothelialized in vivo.146 Additionally,
cells may be subjected to ischemic preconditioning prior to
seeding in vivo to tolerate the new environment better.5,147,148
Alternatively,
current
research
is
focusing
on
pre-
vascularization, in which a capillary bed is built in vitro prior
to implantation using a combination of the elements described
previously. Once implanted, the capillaries on the periphery
of the construct must inosculate with the recipient’s vascula­
ture to deliver oxygen and nutrients to the entire block (see
Fig. 16.13). This approach, although not immediate, is seem­
ingly faster and more reliable than waiting for the host’s capil­
laries to invade the whole tissue from the periphery to its core,
but still does not solve the issue of transient ischemia, which
is not tolerated by a number of tissues. Even though some
studies have achieved a significantly faster inosculation
between 1 to 4 days,108,109,149 further research is still needed to
establish this model of extrinsic vascularization as a feasible
strategy for the engineering of sizeable 3D tissues. Finally,
adequate extracellular matrix (ECM), it is equally important
that the turnover between scaffold degradation and new ECM
deposition by the cells occurs at a rate that neither impedes
the growth of the network by being too slow, nor renders the
construct unstable by being too fast.121
Growth factors
A complete and detailed explanation of the growth factors
and signaling pathways involved in the process of angiogen­
esis is beyond the scope of this chapter. Herein, a general
overview of the main growth factors that participate in the
formation and sprouting of new blood vessels is presented.
Vascular endothelial growth factor A (VEGF-A) is the most
important growth factor involved in angiogenesis in both
health and disease,124 acting mainly by binding to VEGF
receptor 1 (VEFGR-1) expressed in endothelial cells, hemato­
poietic stem cells and inflammatory cells, and VEGFR-2,
which is expressed mainly in endothelial cells. Additionally,
neuropilins 1 and 2 (NRP1 and NRP2) are VEGF-A co-receptors,
which can act either promoting VEGFR-2 activity or indepen­
dently.101 Interestingly, the effect of VEGF-A on the developing
vasculature is dependent on its isoforms and the way it is
delivered. Thus, while soluble isoforms promote vessel
enlargement, matrix-bound isoforms stimulate branching.101
Likewise, paracrine VEGF-A released by tumor, myeloid or
other stromal cells increases vessel branching and renders
tumor vessels abnormal retarding its progression,125 whereas
autocrine VEGF-A, released by endothelial cells, maintains
vascular homeostasis.126 As important as it is in the organism,
VEGF-A has also been suggested to play a critical role in the
design of clinically relevant tissue regeneration graft substi­
tutes through its angiogenic effects and ability to chemoattract
mesenchymal stem cells.127
The fibroblast growth factor (FGF) family comprises 18
members that regulate a number of biological and develop­
mental processes.128 FGF-2 (also known as basic fibroblast
growth factor, bFGF) is the one mostly involved in angiogen­
esis and is produced by a number of differentiated cells such
as keratinocytes, mast cells, fibroblasts, endothelial cells and
smooth muscle cells,129 as well as by adult mesenchymal stem
cells derived from bone marrow, adipose, and dermal tissue.130
FGF-2 has been shown to exert a number of important actions
related to neovessel formation such as stimulating migration
and proliferation of endothelial cells in vivo and encouraging
mitogenesis of smooth muscle cells and fibroblasts, which
induces the development of large collateral vessels with
adventitia.131,132 FGF-2 has been used with success in the pre­
vascularization of scaffolds, with one study showing faster
inosculation of the scaffold in vivo when this factor was
added.109
Platelet-derived growth factor (PDGF) and in particular
PDGF-B released by endothelial cells also participate in the
angiogenesis process mainly by attracting pericytes that will
subsequently provide stability and structural support to the
newly formed vessel.133,134
Other proangiogenic factors identified are placenta-derived
growth factor (PlGF), hepatocyte growth factor (HGF) and
insulin growth factor 1 (IGF). Whilst these factors may play
important roles in developmental and adult angiogenesis, for
prevascularization purposes, it is likely that their role can be
replaced by some of the other agents listed earlier, given the
redundancy of factors and their functions.
BOX 16.3  Animal models for extrinsic vascularization in tissue
engineering
Subcutaneous implantation
Subcutaneous implantation in polydimethysiloxane wells107
Dorsal skinfold chamber335
Cranial window335

246
CHAPTER 16 • Tissue engineering
by Walton et al.160 (Figs. 16.21 & 16.22). We have studied the
mechanism of angiogenic invasion into Matrigel®, as has
been done also by Tigges et al. in mice.161 This highlights
the pre-eminent role of bone marrow-derived macrophages
and pericytes. Macrophages burrow into the matrix, creating
channels which are lined initially by pericytes. Only later do
endothelial cells follow and line the hollow tubes. Vessel-like
networks form initially without any endothelial cell contribu­
tion but these cannot form without macrophages. This elegant
exposé of the process of vascularization was specific to
Matrigel® and may not necessarily reflect the vascularization
process of wounds comprising other matrices. This intrinsic
system of vascularization can support seeded cells and the
histological pattern of their survival clearly demonstrates the
intimate role of the neovasculature in nourishing the cells.
Fibroblasts158 as well as specialized cells including neonatal
cardiomyocytes162 (Fig. 16.23), islet cells,163 and myoblasts5
survived in the chambers but quantification of survival is
difficult and almost certainly many cells apoptose. Delayed
seeding of cells into a prevascularized chamber enhances cell
survival in the rat5 and mouse chambers.163 When ASCs were
implanted into the mouse chamber model in Matrigel® and
traced by sex mismatching or human to animal stains, new
fat tissue formed but the stem cells were found to contribute
only to new blood vessel formation. New fat was of host
origin, indicating that the stem cells did not differentiate into
fat but were inductive for endogenous fat formation. It is
well recognized that MSCs have paracrine effects, including
angiogenesis, cell recruitment, homing, anti-inflammatory,
and antiapoptotic effects.164 We have compared survival
of liver precursor cells cultured on 2D plastic delivered as
cell suspensions with 3D gel-cultured spheroids.165 In the
spheroid form greater numbers of cells implant and survive
in the chamber, suggesting that survival is related to other
factors such as cell–cell microenvironment apart from pure
vascularization. When solid tissue pieces such as muscle were
implanted into the chamber they did not survive but regularly
the dead tissue induced endogenous fat replacement.166
Khouri et al.167 designed a tissue-engineering chamber
(BRAVA®) to induce fat growth for breast tissue replacement.
It consists of a suction device applied to the external surface
of the breast which, when worn regularly for prolonged
periods, pulls the tissue away from the chest wall, creating
tissue injury, inflammation, and a potential space. These are
the key elements of our tissue engineering chamber where
inflammatory cytokines stimulate angiogenesis and attract
stem cells while the vacuum creates the space. Biomechanical
forces including fluid shifts also influence cell behavior.
Subsequently, Khouri included massive volumes of fat injec­
tion which are seen to survive most probably thanks to the
well-vascularized environment generated by the previously
applied suction.168,169 This modification is akin to our
model of seeding cells into a preconditioned prevascularized
chamber. In another original concept of intrinsic vasculariza­
tion for tissue engineering using “autologous microcirculatory
beds”, Chang et al.170 devised an ex vivo perfusion chamber
where an explanted flap of tissue, which is in essence a pre­
formed vascular bed, can be continuously perfused with
oxygen-carrying nutrient fluids for periods of up to 24 hours
through its vascular pedicle. During this time, the flap was
infused with cells intravascularly, including stem cells, which
egressed from the microcirculation into the tissues to form cell
despite the remarkable advances in the area of prevasculariza­
tion that make us believe that the creation of thick, clinically
relevant organs or tissues is within reasonable reach, there is
yet an important hurdle to overcome, namely the fact that the
vast majority of experiments are performed in healthy animals
under relatively controlled conditions in comparison to the
clinical scenario. In the latter, patients with damaged and
scarred tissues, medical comorbidities, tobacco use and other
adverse conditions are the substrate into which these prevas­
cularized scaffolds will be eventually applied and expected to
survive.
Intrinsic vascularization models and in vivo bioreactors
Prefabrication of flaps, where a blood vessel pedicle is
implanted into living tissues to vascularize a dedicated
territory such as skin,150,151 or prelamination, where several
tissues are grafted around the pedicle so that they can be
subsequently transferred as a composite tissue graft,152 has
given tissue engineers insights into the importance of blood
supply and how such vascularization might be incorporated
into tissue-engineered products grown in the body.153 This
process involves the establishment of an intrinsic vasculariza­
tion that grows in tandem with the tissue-engineered product
and nourishes it progressively. When a vascular pedicle or
loop is tunneled into a tissue plane, the surgical trauma,
inflammation, and hypoxia stimulate release of proangiogenic
growth factors and promote angiogenic sprouting, which
links to the pre-existing vasculature (Fig. 16.14). We have
exploited this concept for tissue engineering in a rat model,
where a vessel loop is directed into a protected space that
cannot collapse, such as a chamber box (Figs. 16.15–16.18).154
The vessel loop sprouts new capillaries due to continuing
ischemia within the chamber and shear forces in the loop
further promote angiogenesis. The “wound” cannot contract
or shut down, prolonging the ischemic stimulus, and altered
mechanical forces and fluid shifts influence cell migration and
proliferation within the chamber. The inflammatory environ­
ment of the chamber attracts neighboring and distant cells
to participate in this frustrated repair process in an attempt
to close the dead space, which is progressively replaced
with new vascularized predominantly fibroblastic tissue
(see Fig. 16.18). As the tissue grows, the angiogenic front
moves centripetally and corresponds to the retreating zone
of ischemia.155 Tissue growth is enhanced when a collagen
matrix is included in the chamber and the angiogenic tissue
will readily accept a skin graft, thereby creating a vascular­
ized flap. The angiogenic stimulus continues for at least 20
weeks but is maximal between 7 and 10 days (Fig. 16.19).156
Simcock et al.38 injected human angioblasts intravenously into
rats with an arteriovenous loop chamber and observed that
they could home to the chamber site. This was enhanced by
SDF-1 delivery directly into the chamber and implies that
vasculogenesis plays a role in the vascularization process
(Fig. 16.20). The contribution of vasculogenesis to skin graft
and flap healing has been elegantly investigated by Capla
et al.157 If fibroblasts are added to the chamber at the time
of creation of the loop they survive and can proliferate.158
We have developed a similar model of tissue engineering
in the mouse using a “flowthrough” epigastric artery and
vein in the groin surrounded by a sealed silicone tube, which
includes Matrigel® matrix, a mouse sarcoma extract composed
largely of laminin, and FGF,159 which has also been described

247
Components of tissue engineering
Fig. 16.14  Prefabricated flap in a rabbit thigh. (A) Implantation of a vascular pedicle into the subcutaneous
layer of the rabbit thigh. (B) Injection study at 1 week. (C) Injection study at 3 weeks. (D) Close-up capillary
budding. (Courtesy of M. Hickey, O’Brien Institute.)
A
B
C
D
clusters. The flap was then reimplanted by microvascular
anastomosis into the animal. Similarly, Sekine et al.171 have
also harvested a capillary bed using a piece of rat femoral
muscle tissue attached to the femoral artery and vein. After
connecting it to a bioreactor and perfusing it, the authors
placed cell sheets containing cardiac and endothelial cells and
demonstrated functional connections between the sheet and
the bed. These authors also found that the co-culture of endo­
thelial and cardiac cells as well as the addition of FGF-2 led
to a more robust vascularization of the tissue in the bioreactor.
Finally, they were able to transplant the whole construct back
into the rat as a microsurgical free flap using the carotid artery
and jugular vein as recipient vessels. This model points to a
new direction in organ regeneration or cell transfection to
deliver deficient proteins. Furthermore, it is indeed attractive
in the sense that it circumvents the need to “artificially” create
a vascular network and ensures immediate vascularization
after pedicle anastomosis without having to rely on invasion
or inosculation. Moreover, at a first glance this method seems
applicable to soft tissue reconstruction by harvesting a pedicle
with a small amount of tissue which is expanded in vitro and
transplanted back into the patient; however, it unfortunately
still involves the sacrifice of a vascular pedicle with its inher­
ent donor site morbidity.

248
CHAPTER 16 • Tissue engineering
Fig. 16.19  Angiogenic sprouting from the vessel loop at 10 days. (Courtesy of Dr.
Z. Lokmic.)
Fig. 16.15  Polycarbonate chamber (left) comprising a base plate and a lid (right).
Fig. 16.16  Chamber box with arteriovenous loop.
Transparent
plastic cylindrical
chamber with lid
Hole for insertion
of vessel loop
Anteriovenous
fistula
Fig. 16.17  Chamber implanted (A) in rat groin, where (B) it is closed and buried.
A
B
Fig. 16.18  Chamber opened after 6 weeks showing tissue formation.

249
Components of tissue engineering
environment through focal adhesion transmembrane path­
ways as discussed earlier,41 stimulating cell migration and
growth. However, quantification of this effect is difficult in
vivo and the precise process remains unclear.
Encouraged by the promising findings in animals, we
questioned whether the tissue-engineering chamber might
have a similar effect in the clinical setting and thus we scaled
up our research into a human trial.174 Five patients were
recruited, all candidates for breast reconstruction who wished
to avoid breast implants if possible and were unsuitable for
autologous reconstruction. A pedicled fat flap based on the
thoracodorsal system ranging from approximately 5 to 50 mL
was harvested, transposed onto the chest wall and placed
inside a perforated, acrylic, custom-made chamber ranging
from 180 to 360 mL. Magnetic resonance imaging was per­
formed every three months and the chambers removed six
months after implantation. One patient was withdrawn from
the trial at 7 weeks due to painful rubbing of the chamber’s
rim on her ribs. In another patient there was significant expan­
sion of the flap by 6 months on MRI and it was mutually
agreed to leave the chamber in situ for twelve months total.
At exploration the whole chamber space had filled with fibro-
fatty tissue surrounded by a fibrous capsule representing an
In a different model of intrinsic 3D tissue engineering, a
vascularized flap of fat tissue, when implanted into an empty
chamber space, will expand and attempt to fill the space. We
initially observed in our mouse Matrigel® model that if a fat
pad was used to plug one end of the chamber tube and wax
the other, fat gradually replaced the Matrigel® by direct exten­
sion from the pad.159 When we placed an isolated fat pad on
a vascular pedicle into the chamber in a rat even without
matrix, the fat tissue expanded. When perforations were
included in the chamber shell the space completely filled with
tissue, approximately 40% of which was new fat.172 This result
was repeated in a pig model which grew over 50 mL of new
fat (Fig. 16.24).173 Although surgical trauma, ischemia, and
inflammation play a role through biochemical signaling in
this new tissue growth, it is likely that the tissue flap finding
itself in an empty space reacts to the changed biomechanical
Fig. 16.20  Human angioblast, DiI-labeled, homing to the rat chamber after
intravenous injection. (Courtesy of J. Simcock and G. Mitchell, O’Brien Institute.)
Fig. 16.21  Tissue-engineering silicone tube chamber in a mouse. (Reproduced
from Cronin KJ, Messina A, Knight KR, et al. New murine model of spontaneous
autologous tissue engineering, combining an arteriovenous pedicle with matrix
materials. Plast Reconstr Surg. 2004;113:260–269.)
Fig. 16.22  Vascularization into the Matrigel® in a mouse chamber at 4 weeks.
Fig. 16.23  Neonatal cardiomyocytes seeded into rat chamber showing maximum
survival around the vessel loop. (Courtesy of A. Morritt, O’Brien Institute.)

250
CHAPTER 16 • Tissue engineering
Fig. 16.24  (A) Chambers in pig model, 80 mL volume. (B) 5 mL volume fat flap
on epigastric pedicle inserted into the chamber. (C) Buried for 6 weeks.
(D) Removal at 6 weeks, filling the chamber. (E) Approximately 50% volume is new
adipose tissue.
A
B
C
D
E
expansion of the flap from approximately 30 mL to 210 mL
(Fig. 16.25A,B). Conversely, the flaps in the remaining three
patients failed to expand and instead were encased in a thick
fibrous capsule. Although the trial was unsuccessful, the
unquestionable finding of stable tissue growth to the cham­
ber’s capacity in one of the patients, which was stable at 6
months after chamber removal (Fig. 16.25C), provides a proof-
of-concept that clinically relevant volumes of autologous
tissue can be generated using the tissue-engineering chamber
in humans. The difference in the consistency of results within
this trial and in comparison to those obtained in animals
warrants further investigation.

251
Emerging technologies
when biological components are included. Whilst this tech­
nology has been known for processing plastics for over 30
years, major improvements in the last decade have led to
simultaneous reductions in the costs involved and dramatic
increases in flexibility. Various types of printing equipment
can now be used to produce three-dimensional objects from
computer-aided designs (CAD) simply and affordably, as
illustrated in Fig. 16.26. At its simplest, a 3D printer resembles
a hot-glue gun or syringe delivery system that lays down
droplets or fibers of a chosen material in 3D patterns built up
Emerging technologies
A range of engineering and technology developments that
may significantly impact tissue engineering have emerged
over recent years.
3D printing
The most striking development is additive manufacturing via
3D printing, sometimes termed bioprinting or bioplotting
Fig. 16.25  (A) Intraoperative view upon chamber removal. The space is completely filled with fibro-fatty tissue surrounded by a fibrous capsule. Original flap volume was
30 mL and chamber space 210 mL. (B) Detail of tissue generated inside the chamber showing adipose tissue under fibrous capsule. (C) Patient at six months after
chamber removal. Newly formed tissue remained stable and softened to a “fat-like” consistency. (From Morrison WA, Marre D, Grinsell D, et al. Engineering of adipose
tissue in humans using the tissue-engineering chamber: a pilot study. EBioMedicine. 2016; in press.)
A
B
C
Fig. 16.26  Methods of producing three-dimensional (3D) biomaterial constructs via solid freeform fabrication. (A) Stereolithography. (B) Selective laser sintering. (C) 3D
printing. (D) Wax printing. (E) Fused deposition modeling. (F) Bioplotter. (A–E, From The Worldwide Guide to Rapid Prototyping website © Copyright Castle Island Co. All
rights reserved. F, From Hollister SJ. Porous scaffold design for tissue engineering. Nat Mater. 2005;4:518.)
Laser
Laser
Scanner
system
Powder
bed
Roller
Powder
bed
Liquid
adhesive
supply
Particle
collector
Milling
head
Ink-jet
head
Roller
Fabrication
piston
Object
being
fabricated
Vat
Photopolymer
Moveable
table
Powder-
delivery
system
Fabrication
piston
Plastic filament
supply coil
Table
Extrusion
nozzle
Sterile
filter
Thermostat
Sterile
compressed air
Ultraviolet
lamp for
disinfection
Sterile environment
(laminar flow)
Plotting material
(with cells)
Plotting
medium
3D objects
(with cells)
Object and
support
materials
x-y
stage
x-y-z
stage
Powder-
delivery
system
Scanner
system
A
B
C
D
E
F

252
CHAPTER 16 • Tissue engineering
In vitro bioreactors
Cell culture in 3D constructs presents challenges in delivery
and removal of nutrients and waste products in the absence
of a blood supply.181 The delivery of oxygen is challenging due
to its relatively low solubility in cell culture media and the
time required for diffusion to occur into human-scale 3D
constructs. Delivery of macromolecules such as growth factors
also presents challenges due to their large size and thus rela­
tively slow diffusion. In addition, mechanical forces, hydro­
dynamic conditions and electrical stimulation can significantly
affect cellular processes and tissue development. Therefore,
advanced bioreactors may be helpful to control the microen­
vironment of cells in vitro to encourage the desired cellular
processes and tissue development whilst also optimizing the
mass transport of oxygen, nutrients and waste products to
allow 3D tissue constructs to be developed.182
Computational modeling
Mathematical modeling can assist in design of tissue-
engineering constructs and processes, providing insights into
factors governing tissue growth. For example, different cell-
seeding patterns and timing can be investigated using a
mathematical model of a tissue-engineering scaffold during
vascularization in vivo to provide insights into oxygen con­
centration distributions and cell migration rates. Such models
are necessarily simplifications of the complexity of the bio­
logical processes at play but may aid experimental design and
interpretation and so help reduce trial-and-error experimenta­
tion and animal use.183,184
Testing and characterization of tissue
engineering approaches
Characterization of new or modified biomaterials, cells, and
methods for tissue engineering is critical to their potential
pathway to clinical trials and ultimately to clinical and com­
mercial success. Systematic testing is not only a requirement
of regulatory and ethical approval during the development
process, but also crucial to developing understanding and
thus being able to improve tissue-engineering methods
further. Testing should aim to allow prediction of the long-
term safety and efficacy of the tissue-engineering approach
under consideration. Ideally, in vitro testing would be used to
screen tissue-engineering constructs and methods in order to
minimize animal testing for ethical reasons. However, it
remains a significant challenge to predict in vivo responses
from in vitro tests and so significant animal testing is normally
required before clinical trials.
Standardized testing methods are highly desirable so
that objective comparisons can be made between different
approaches and construct components across different labo­
ratories. Testing should be undertaken under conditions
simulating the targeted tissue-engineering application, rather
than relying on test results for the same material in different
applications where phenomena may differ significantly.185
Physicochemical characterization of biomaterials and con­
structs made from biomaterials should be undertaken to
ensure the targeted composition, absence of significant resi­
dues or impurities, and desired morphology are achieved.
layer by layer as the printer nozzle and/or printing stage
move according to the CAD requirements. A wide variety of
materials can be processed using different kinds of 3D print­
ers, including polymers, hydrogels, metals, ceramics and even
living cells. Many detailed reviews have been published on
aspects of this technology and its potential in tissue engineer­
ing in recent years.88,175
Many 3D printers and other forms of rapid-prototyping
equipment are limited to a certain category of raw materials,
e.g., photo-polymerizable or melt-processable polymers with
specific physical properties. However, this limitation is also
being overcome with printers such as the 3D-Bioplotters
(EnvisionTEC GmbH), which are able to print in numerous
user-developed materials including melt-processed poly­
mers, hydrogels and viable cells.175,176 It is also possible in
many devices to print multicomponent mixtures in defined
patterns to create more complex objects. Challenges remain­
ing to be resolved with this rapidly evolving technology
include the potential of damage to cells and biomolecules
during printing due to shear forces in the nozzles, par­
ticularly when high spatial resolution is desired. Different
printing methods have limitations in material choices and
printed construct properties such as pore and fiber size
ranges, construct surface properties (e.g., roughness) that
can affect cell–material interactions, as well as accuracy,
processing cost and speed. Appropriate quality control,
predictability of properties such as construct mechanical
strength and defect frequencies, the ability to print in a
diverse range of medical grade materials, as well as regula­
tory and industry guidance are areas for further improve­
ment before this technology can be widely adopted in tissue
engineering.177
3D printed polymer scaffolds have been used clinically in
a range of applications, mainly for hard tissue repair, typically
by creating a customized implant using medical imaging data
for a particular tissue defect in a patient. Notable examples
include scaffolds for craniofacial bone repair and some recent
reports on airway splints for tracheobronchomalacia.95,177 Such
biodegradable scaffolds or implants are sometimes termed
“4D” as they change with time, ideally resorbing as the
patient’s tissues develop.
Another area of rapidly increasing activity is production of
tissue- or organ-on-a-chip modules often created using micro­
fluidic devices and 3D printing. Such constructs rely on tissue
engineering principles and provide opportunities to study
tissue development, diseases and screen drugs in more real­
istic in vitro environments.178,179
While biofabrication via 3D printing facilitates the manu­
facture of custom-designed tissue engineering constructs
with excellent spatial control of the initial distribution cells
and biomaterials, the major challenge of achieving timely
vascularization in 3D tissue engineering constructs discussed
above for constructs made in other ways, still applies for
3D-printed constructs. This is often neglected in reports about
the potential of this technique, especially in the media where
hype sometimes overtakes reality. Research to address this
problem has included inclusion of interconnected channels
in 3D-printed constructs to allow perfusion in vitro and
in a recent development, Kolesky et al. used a custom-
made bioprinter to fabricate heterogeneous constructs that
included multiple cell types and vascular channels lined with
endothelial cells.180

253
Case studies and status of tissue engineering for specific tissues
However, graft durability and functionality is often poor, with
wound contraction, scar formation, and impaired sensation
being frequent complications in addition to graft failure. To
improve this, skin tissue engineering has focused on generat­
ing a dermal component in addition to the epidermal layer.
Such tissue-engineered skin substitutes would function
beyond simple wound coverage and barrier protection. They
could be applied in the treatment of chronic wounds with
impaired wound healing such as diabetic ulcers, in skin
disease models for example with vitiligo and psoriasis, and
for drug and cosmetic testing. Additionally, oral mucosa
tissue engineering draws many similarities with skin,192 and
this will no doubt be useful in craniofacial reconstructions.
The principles of skin tissue engineering follow the same
concept as other tissues and involve cells, scaffold, and vas­
cularization. Cell sources include fibroblasts to promote col­
lagen and other ECM structural integrity but also to address
the fundamental problem of biofilm that protects bacteria
from attack by neutrophils and renders local fibroblasts
senescent. Live fibroblasts can be incorporated into the
implant to attract and activate neutrophils through the secre­
tion of proinflammatory cytokines and growth factors.193,194
Other cell types include epidermal-derived stem cells,195 and
MSCs especially for wound healing,196 and melanocytes.197
Bulge cells from hair follicles and co-cultures, especially with
neonatal dermal fibroblasts, offer some hope of regeneration
of functional skin elements, including hair follicle and
sebaceous glands.198 Cutaneous nerve regeneration may be
enhanced through the addition of Schwann cells, or epidermal-
derived stem cells and other MSCs which are capable of dif­
ferentiation into both neurons and Schwann cells.199
Scaffolds are a key aspect of engineered skin. Substitutes
initially served as passive matrices for ingrowth from the
wound bed full of blood vessels and dermal elements, but
more recently complex structures have been developed
comprising of biomimetic proteins and material, 3D culture
techniques, and altered surface chemistry and topography to
influence cell adhesion, growth and differentiation, durotaxis,
spatial cueing, and mechanoregulation.200,201 They can also
facilitate nerve regeneration,199 and may include antibacterial
agents such as PLGA films and polylactic acid nanofibers
loaded with silver ions or nanoparticles.202–204
Current products suffer from poor integration to the wound
bed as a result of inadequate vascularization and this impacts
on the survival of implanted cells. Recent efforts to address
this include the use of angiogenic biomaterials and prevascu­
larization of constructs (covered in earlier sections), as well as
the use of stem cells such as adipose-derived stem cells which
abundantly secrete angiogenic growth factors and can also
contribute directly to regeneration by differentiating into
fibroblasts, epithelial, and endothelial cells.205
Adipose tissue engineering
Adipose tissue engineering holds great interests to plastic
surgeons for reconstruction of soft tissue deformities such as
Poland syndrome, Romberg syndromes, aging, breast post­
mastectomy and traumatic deformities, burns, and postirra­
diation defects.
Adipocytes are fragile, prone to ischemic death,206 very
difficult to grow in cell culture, and do not proliferate in vivo.
Consequently, precursor populations derived from bone
Materials which are nominally the same may perform differ­
ently in tissue-engineering applications due to differences
such as the presence of impurities, differences in molecular
weight or polydispersity in polymers, or variations in surface
properties.
Risks, ethics and regulation
Risk management is a critical consideration in clinical and
commercial tissue-engineering efforts. A number of potential
products developed early this century suffered from problems
with regulatory trials and product launches, along with
challenging investment conditions internationally.186 Tissue-
engineering approaches have significant potential benefits to
patients and the health system but also involve risks, which
may include human error (e.g., mix-ups of autologous cells),
biomaterial performance variations from predictions (e.g.,
changes in degradation rates in vivo), quality control problems
(e.g., microbiological contamination), and variations in
patient-specific responses. Most of these can be minimized by
rigorous management and scientific processes (e.g. through
GMP process requirements, process automation and quality
control), whilst patient-specific responses will only be fully
assessed once sufficient clinical trial data have been collected.
Many of the commercial successes in the field of tissue engi­
neering to date have avoided the risks involved in delivering
biological materials, rather using acellular products.186
The ethical considerations related to tissue engineering
have been reviewed187,188 and many of these are common with
other fields of medicine and biomaterials. The sources of
biological material to be used, particularly the potential use
of ES cells, are, unsurprisingly, prime considerations. The
extent of animal testing is a further issue of concern, which
may be partly addressed by effective in vitro testing and
well-designed experimental approaches, as discussed above.
Animal trials should be designed to simulate the final clinical
application as closely as possible to avoid incremental
approaches to the final design with numerous animal trials at
every stage.
The regulatory approvals required for tissue-engineered
products depend on the jurisdiction involved and have
evolved significantly over the last decade around the world.
In the case of the US Food and Drug Administration (FDA),
products may be characterized as tissues, biological products,
medical devices, or combination products and numerous
standards and guides have been developed for particular
categories of products.189,190 The ISO10993 series of standards
addresses “Biological Evaluation of Medical Devices” and
numerous ASTM guides provide details on testing and char­
acterization requirements for different types of tissue engi­
neering products.
Case studies and status of tissue
engineering for specific tissues
Skin tissue engineering
Biological wound dressings have been long-established from
cultured sheets of keratinocytes,191 and together with skin
banks this has saved the lives of many burns patients.

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CHAPTER 16 • Tissue engineering
FACS-sorted or SVF-derived human ASCs or bone marrow
stem cells were implanted, endogenous fat formed.208 If the
fat graft or stem cells are substituted by an inflammagen such
as yeast particle zymosan, fat also forms in Matrigel®.209 Taken
together with the early appearance of macrophages in the
chamber and their critical role in angiogenesis and adipogen­
esis as shown by macrophage ablation studies,210 it suggests
that the signaling mechanism by which fat grafts stimulate fat
growth is an inflammatory one. This is consistent with the
emerging notion that MSCs largely function through their
paracrine effects especially to promote angiogenesis and cell
signaling. The cytokine profile of the mouse chamber contents
clearly shows inflammatory angiogenesis predominating for
the first 7 days, at which time it changes to adipogenesis
with the appearance of peroxisome proliferation-activated
receptor-γ and other markers.211 Whether this event sequence
is peculiar to Matrigel® matrix is not yet clear. Nevertheless,
increasing evidence suggests that initial cell death within fat
grafts is followed by neoadipogenesis, which is mediated by
stem or progenitor cells within the grafted material with the
recruitment of host cells at the implantation site.212,213 This is
supported by recent clinical studies reporting the enhanced
survival of fat grafts enriched with SVF/ASCs when com­
pared to fat grafts alone.214,215 However, whether the fat graft
itself survives at least in part remains unanswered, with a
wide variability of results in the literature.216 Furthermore,
questions remain as to whether the stem cells themselves
survive and contribute directly to new fat, or if they represent
a concentrated dose of inflammatory paracrine factors which
induce endogenous fat production.
Matrix plays a key role in adipogenesis, as does angiogen­
esis. Multiple synthetic and natural scaffolds have been trialed
in vitro and in vivo with and without FGF and/or seeded cells.
Matrigel® is both angiogenic and adipogenic,217 and several
Matrigel® substitutes are under study, including derivatives
of muscle (myogel)55 and adipose-derived matrix (ADM).218
The ultimate clinical product would be an adipogenic inject­
able matrix gel with or without processed lipoaspirates. ADM
has so far shown promise in stimulating the adipogenic dif­
ferentiation of ASCs both in vitro and in vivo,218 and can be
used as an off-the-shelf product to regenerate subcutaneous
fat in soft tissue defects created in the rat back.219
Another important area is the role of mechanical forces
that influence cell behavior, including differentiation. We
believe this is critical to the growth seen in the chamber
space, particularly in those models where there was no matrix
or seeded cells such as the adipose tissue growth seen in
the pig model (see Fig. 16.24) and clinical study (see Fig.
16.25), corroborated more recently by others using a rabbit
model.220
Muscle tissue engineering
Tissue engineering of skeletal muscle has potential in vivo
applications for the treatment of muscular dystrophy, muscle
loss from injury and post-tumor resection, chronic denerva­
tion, and targeted delivery of proteins and drugs by gene
transfection of muscle cells. Ex vivo applications could include
models for the study of exercise physiology or muscle diseases
or drug screening. It could also in theory find a commercial
use for tissue-engineered meat production from animal stem
cell sources.221
marrow or fat SVF and sorted by adherent cell culture or
fluorescence-activated cell sorting (FACS) are the preferred
cell source for the traditional loaded-scaffold paradigm of
tissue engineering. However, this direction has been singu­
larly unsuccessful for fat, no matter what scaffold is chosen,
and most clinically focused research is returning to modifica­
tions of fat injection, possibly with matrix material as a more
likely route to clinic.
Attempts to understand the fate of fat-grafting dates back
almost a century when Marchand207 grafted fat onto the brain
of animals. Upon exploration, he identified two cell types: (1)
fat, which he interpreted as having survived a graft; and (2)
histiocytes, which had come from the hosts and appeared to
be turning into fat, which he labeled as replacement. Since
then, a long debate ensued as to whether cells in grafted
fat survive or signal endogenous fat production. When we
inserted human fat grafts into our mouse chamber tissue-
engineering model containing Matrigel® and FGF-2, new fat
formed but it was of mouse origin (Fig. 16.27). Similarly, when
Fig. 16.27  (A) Spontaneous endogenous fat formation in Matrigel® in a mouse
model of tissue engineering after insertion of a small nonvascularized fat allograft.
(B) Normal fat tissue grown in the chamber. Note the extensive vascularization
sprouting from the vascular pedicle. (Courtesy of F. Stillaert, O’Brien Institute.)
A
B

255
Case studies and status of tissue engineering for specific tissues
include the manipulation of the cell culture surface topogra­
phy using nano/micro-patterning techniques, ECM printing,
as well as the application of mechanical strain and electro­
magnetic fields in order to generate myotubes and skeletal
muscle sheets.246
ECM compositions play a key role in the alignment and
differentiation of myoblasts. Hydrogels facilitate self-assembly
of suspended cells into transplantable constructs, and
collagen,247 fibrin,248 alginate,249 hyaluronic acid,250 PEG,251
and extracellular matrix derivatives such as Matrigel®252 and
myogel253 have been used. Various synthetic matrices, includ­
ing PLGA mesh,254 PCL,255 chitosan,93 glass fibers,256 electro­
spun nanofibers,257,258 and wavy silicone surfaces259 influence
regulation of cell alignment differentiation and proliferation
but the elasticity, toxicity, and biodegradability are critical to
in vivo application. Decellularized biological scaffolds may be
more physiologically attuned with this regard, as they contain
relevant ECM components in their relative compositions, are
non-immunogenic, have a preformed shape, and likely retain
biologically active cytokines and growth factors and func­
tional cues. Decellularized scaffolds from porcine urinary
bladder (MatriStem™) have been used to stimulate local
muscle regeneration in patients with volumetric muscle loss
following severe leg injury.260 By placing these scaffolds within
the injured muscle close to viable tissue followed by intensive
physical therapy, biopsy at 6–8 months showed new muscle
tissue with resorption of the scaffold, correlated with CT/MRI
imaging. Improved functional outcomes were reported in 3
out of 5 patients despite an original muscle loss ranging from
58–90%. It is expected this could be further optimized through
the addition of cells to the scaffold, such as in a study where
decellularized skeletal muscle was reseeded with myoblasts,
fibroblasts, perivascular stem cells, and endothelial cells and
used to reconstruct a rat abdominal wall defect.261 This
approach has also been taken to reconstruct a bioartificial limb
from decellularized rat and primate forearms.262 Scaffolding
from the soft tissue including muscle was recellularized in a
bioreactor with endothelial cells, myoblasts, and fibroblasts
followed by electrical stimulation and coverage of the forearm
with a skin graft (Fig. 16.28). Despite no integration of neural
tissue, the rat paw was seen to clench and unclench in response
to electrical stimuli with approximately 80% the strength of a
newborn rat. Functional anastomosis of blood vessels could
be achieved upon transplantation of these forearms with
preserved response to electrical stimuli. Although this remark­
able proof-of-concept study still requires many steps to
become clinically relevant, it provides encouragement in the
field of decellularized scaffolds.
Vascularization remains a large hurdle in the upscaling of
engineered tissue in vitro, and the survival of transplanted
tissue. In recognition of this, in vitro approaches may combine
myogenic cells with endothelial cells within 3D scaffolds,263,264
or combine the delivery of growth factors such as VEGF or
insulin-like growth factor 1 together with cells in a scaf­
fold.247,265 Scaffolds themselves can be tailored to have sus­
tained release of these growth factors to mitigate their short
half-life, and encourage blood vessel ingrowth from the host
into the engineered tissue upon transplantation.266 Applying
intrinsic vascularization is another option, and we have used
both the mouse flow-through pedicle model and rat arterio­
venous loop model to demonstrate the survival of implanted
myoblasts.252,267 Cell survival can be enhanced through the use
Muscle is unique functionally with its densely packed
parallel-aligned multinucleated muscle fibers which contract
axially through calcium-sensitive actin and myosin filaments.
This in turn involves neural connections functionally linked
through acetylcholine release and its receptors. Addressing
this specialized architecture with the incorporation of vascular
and somato-sensory elements to create fully innervated, vas­
cularized skeletal tissue that can be transplanted and function
in vivo remains a major challenge.
Cells used for muscle tissue engineering include myoblasts,
the progenitor cell to mature myocytes which can be isolated
from skeletal muscle and expanded in vitro. However, differ­
entiation in 2D culture after extended passaging is difficult,
and requires ECM coating or 3D culture to facilitate myoblast
alignment and differentiation.222,223 Another cell source
includes the more immature satellite cells, which are resident
stem cells in skeletal muscle residing below the basal lamina
of myofibers. They represent 2–7% of the cell population in
adult muscle and are capable of self-assembly into new myo­
tubes following injury and stimulation.224 However, expansion
of these cells after isolation is problematic, with reduced
proliferation and regenerative capacity with prolonged
culture.225 Satellite cells can be replenished by pericytes, which
are found in the walls of small blood vessels.226 Pericytes
isolated from a wide variety of tissue including skeletal
muscle are also capable of myogenesis both in vitro and in
vivo, regardless of tissue origin.227 They are closely related to
mesoangioblasts, which are perivascular cells found in larger
vessels and are also capable of contributing to muscle regen­
eration.228 Both pericytes and mesoangioblasts are phenotypi­
cally similar to MSCs, which are also attractive candidates.
MSCs can be derived from a wide variety of tissues, as dis­
cussed above, home into sites of injury, secrete useful paracrine
factors, and directly contribute to new muscle.228 More recently
studies have also focused on the generation of myoblasts,
myogenic MSCs, and mesoangioblasts from pluripotent stem
cells such as iPS or ES cells.229–232 This opens a gateway for the
use of autologous cells for tissue engineering, as well as the
transplantation of genetically corrected cells for conditions
such as muscular dystrophy.229
Strategies to improve the myogenic differentiation of cells
include the use of 3D perfusion bioreactor systems233 and
co-culture of myoblasts with various other cells such as SVF
(containing ASCs)233 and fibroblasts.234 Co-culture with neural
tissue in a 3D matrix can also enhance in vitro muscle func­
tion,235 as well as generate neuromuscular junctions with
functioning acetylcholine receptors.236 Electrical stimulation
of myoblasts in 3D collagen culture mimics nerve stimulation
and promotes differentiation to myotubes and ECM deposi­
tion.237 Mechanical stretch or force organizes myoblasts into
functionally aligned myotubes and favorably influences fiber
diameter and cell number238,239 by myofilament gene regula­
tion and protein expression.240,241 In vitro preconditioning of
myoblasts by cyclic strain in a bioreactor improved the con­
tractility of engineered tissue.242 Other specialized culture
systems include the co-culture of myoblasts with fibroblasts
on Seran wrap or laminin-coated plates stretched out between
pins. This led myoblasts to orientate themselves axially and
separate from the wrap upon differentiation to assemble into
a cylindrical core of contractile muscle using the pins as
tendon-like anchors, which were termed as myoids.243–245
More recently, sophisticated approaches to cell alignment

256
CHAPTER 16 • Tissue engineering
factors can also increase the success of tissue-engineering
strategies for nerves.268,276 Three-dimensional printing could
be used in the fabrication of these nerve guidance conduits,
and there is potential to integrate this technology with bionic
limbs by connecting bioengineered scaffolds as nerve guides
to electrodes in the prosthetic limb.
The environment and cellular responses to injury in the
central nervous system tend to inhibit regeneration. Stem cell
therapies for spinal cord injury including several human trials
have shown some regenerative responses, but the mechanisms
of how they may repair the spinal cord are not well under­
stood.277 Direct contribution to new tissue, release of trophic
factors, and modulation of inflammatory cytokines particu­
larly in the early stages of injury have been postulated. Tissue
engineering could be applied in this area to enhance the
delivery of regenerative cells, scaffolds, and neurotrophic
factors to promote axonal regrowth following injury. Scaffolds
or hydrogels can aid cell adhesion, retention and differentia­
tion, or be used to deliver growth factors and drug molecules,
and oriented pores and channels can guide axonal growth.278
Interestingly, the use of an acellular scaffold may forestall
immediate scarring of the spinal cord following injury, pro­
viding a more permissive regenerative environment.279
Tissue engineering of blood vessels
A number of strategies to tissue engineer blood vessels in vitro
or in vivo as an alternative to autografts, allografts, or nonde­
gradable polymer prostheses have been developed and tested
in animal models and in the clinic with promising results.280
The engineering of capillary networks in scaffolds has been
covered earlier in the chapter, but advances have also been
made in the development of larger macrovessels. Collagen-
based scaffolds containing a polyester mesh are available
commercially (Omniflow II from Bio Nova International) and
are used clinically as a biosynthetic vascular graft for hemodi­
alysis access.281 Decellularized veins have been used as
allografts, although a high rate of aneurysmal and thrombotic
complications have been noted.282 Grafts have also been made
by culturing allogeneic smooth muscle cells on PGA scaffolds
to encourage ECM deposition and remodeling of the synthetic
material, followed by detergent-based removal of these cells to
retain only the ECM skeleton.283 This Humacyte technology is
of techniques such as in vitro preconditioning to stimulate
pro-survival signaling pathways (Akt and ERK1/2), or in vivo
prevascularization with delayed seeding to prepare a dense
capillary bed prior to cell delivery. However, to generate
larger volumes of vascularized three-dimensional skeletal
muscle it is likely that a combination of vascularization strate­
gies would be required during both the in vitro and in vivo
phase, with amalgamation of advances in cell co-culture,
growth factor delivery, scaffold tailoring and intrinsic vascu­
larization. This will need to be accompanied by appropriate
electrical stimulation and neuronal integration to maintain the
growing muscle tissue.
Tissue engineering of nerves
Nerves in different locations have distinct microenvironments
and thus the optimal tissue-engineering strategies for each
will differ. Tissue-engineering strategies may be used for
guided nerve regeneration using, for example, degradable
biomaterial tubes or scaffolds. There are now a range of
FDA/Conformit Europe approved neural scaffolds avail­
able which largely consist of collagen I, but include PVA,
PCL and decellularized matrix derivatives.268 Hollow tubes
can function effectively as nerve guides for gaps of up to
10 mm in the peripheral nervous system under a range of
clinical situations. For example, tubular collagen nerve guides
(NeuraGen®, Integra LifeSciences) are being used clinically to
treat peripheral nerve injuries.269 The critical gap length that
may be successfully treated using nerve guides can be longer
than 10 mm in primates and can be further increased by addi­
tion of cells and cell supports such as biomaterial fibers or
hydrogels.270 It has been shown that use of oriented hydrogel
matrices (e.g., fibrin or collagen) within nerve guides results
in a larger critical gap length than use of the same matrices
without alignment. Electrospun fibers can provide mechani­
cal support to guide nerve regeneration, depending on their
diameter, alignment, and spacing, and nanofibers may also
be developed within a hydrogel.271 Including conductive
materials such as gold nanoparticles or carbon nanotubes can
facilitate the use of electrical stimulation, which can increase
neurite outgrowth and maturation.272,273 Schwann cell grafts
including those generated from iPSCs,274,275 and the addition of
controlled-release vehicles containing drugs or neurotrophic
Fig. 16.28  Recellularization of decellularized rat arm scaffolds. (A) Progressive decellularization of whole rat arms. (B) Recellularization of arm scaffold. (A, Reprinted from
Jank BJ, Xiong L, Moser PT, et al. Engineered composite tissue as a bioartificial limb graft. Biomaterials. 2015;61:246–56, with permission from Elsevier. B, Reproduced with
permission from Dr. Harald C. Ott and Dr. Bernhard J. Jank, Massachusetts General Hospital and Harvard Medical School.)
A
B

257
Case studies and status of tissue engineering for specific tissues
Scaffolds which are osteoinductive have also been produced
and shown to enhance bone growth. A wide range of natural
biopolymers have been studied for bone tissue engineering,
including collagen, gelatin, silk fibroin, chitosan, alginate, and
hyaluronic acid.299 Among inorganic compounds, bioactive
glasses composed of silica-containing calcium, sodium and
phosphorous oxides have been termed “bioactive” because
they can form strong bonds to hard and soft tissues in vivo.300
These materials have been marketed commercially as Bioglass
and used clinically, primarily in nonload-bearing applications.
Composite materials are attractive as they can be used to
control biodegradation and can overcome brittleness. This
strategy mimics the composite structure of bone, being an
inorganic–organic composite structure. Examples include
nanocomposite structures combining collagen and calcium
phosphate, as well as chitosan–gelatin and ceramic scaf­
folds.299,301 Suspensions or pastes of ceramic or composite
materials may also be used in surgical repair or via injection
to encourage bone growth for small-volume repairs.
In addition to the influence of the scaffold, the osteoinduc­
tive growth factors, bone morphogenic proteins (BMPs), may
be used to enhance bone tissue growth. Other growth factors
(e.g., TGF-β1, FGF-2, platelet-derived growth factor) may also
enhance bone formation, as summarized by van Gaalen et al.300
Tissue engineering utilizing BMPs has progressed to a number
of clinical trials, primarily in spinal surgery, showing some
promising results.92,300 VEGF plays a dual role in angiogenesis
and osteogenesis, and its encapsulation in PLGA scaffolds has
been shown to increase both vascularization and bone forma­
tion.302 However, the direct use of growth factors in a clinical
setting will need to proceed with caution with particular atten­
tion to safety and optimizing delivery systems. Using cells that
secrete osteoinductive and angiogenic growth factors may be
another option. This includes cells such as MSCs, as well as
endothelial cells, including endothelial progenitor cells, which
have been shown to secrete BMPs and TGF-βs and contribute
to bone healing.303,304 Furthermore, the addition of osteogenic
cells, such as osteoprogenitor cells found in the bone marrow
or periosteum, MSCs, and more recently iPSC-derived MSCs
or osteogenic progenitors, particularly in addition to support­
ive cells and growth factors, may accelerate bone repair through
direct tissue contribution but also ECM deposition and interac­
tion with the host environment.295,296
Cartilage tissue engineering
Tissue engineering of cartilage is relatively well developed as
the problem is simplified by not requiring vascularization
within the developing tissue. Several cell and tissue-
engineering treatments for cartilage repair have been estab­
lished clinically. However, clinical use is not widespread as
yet due to limitations in efficacy, consistency, and applicabil­
ity.305 Autologous chondrocyte implantation for articular car­
tilage repair has been approved by the FDA and implemented
clinically (Carticel®, Genzyme Biosurgery, US).294 Although
this procedure has shown good clinical outcomes, it is labor-
intensive in the cell expansion and multi-stage procedure, and
limited in the defect size that can be treated.306,307
Biomaterial support structures may be used to treat larger
cartilage defects, deliver cells, and provide critical mechanical
support until the tissue is sufficiently developed. A possible
timeline for cartilage repair using a biomaterial implant is
now undergoing clinical trials. More complex scaffold-based
vascular grafts have also been developed, relying on micro-
printing and electrospun nanofibers from natural and synthetic
polymers to fabricate a tri-layered structure mimicking the
layers in large vessels, with increased mechanical strength.284
These approaches rely on the rapid revascularization of the
scaffold by host cells after implantation to form a new blood
vessel. Loading scaffolds with cells prior to implantation has
also been attempted,285,286 and can use various cells other than
endothelial cells including endothelial precursor cells, ASCs,
pericytes, and MSCs.287 Cells can also be incorporated as micro­
tissue units into scaffolds in the form of spheroids.288 However,
the use of cells in macrovessel tissue engineering has focused
more on a scaffold-free approach. This relies on cell sheet
engineering, where sheets of cultured dermal fibroblasts and
their surrounding ECM are rolled into tubes which are then
seeded with the patient’s endothelial cells.289 This technology
has since progressed to clinical trials as the Lifeline tissue-
engineered blood vessel.289–291 In another alternate strategy,
Campbell et al. have developed an intraperitoneal graft model
where silastic tubing is implanted into the peritoneal cavity in
vivo, leading to a foreign body reaction consisting of an inner
collagen layer, middle myofibroblast layer and outer mesothe­
lium which form tubular “vessels” comprised of autologous
cells. These vessels have been successfully grafted into the
circulatory system of the host animal where they were shown
to remodel in response to the new environment and remain
patent for up to 16 months in rabbits.292,293
Bone tissue engineering
Significant research efforts have been made to tissue engineer
both load-bearing and non load-bearing bone to meet the
demands for alternatives to autografts or allografts, which
have their inherent disadvantages and supply limitations.
However, limited translation to clinical application has
occurred to date.294,295 Bone tissue-engineering strategies
generally involve a biomaterial scaffold of suitable mechani­
cal strength with or without added cells and/or biomolecules
to encourage bone growth. Vascularization has often been
neglected in bone tissue engineering, but is increasingly rec­
ognized to play a role beyond simple cell survival and
nutrient/waste exchange, as normal bone repair occurs in
close association with blood vessels and angiogenic factors,
particularly VEGF, play a role in initiating and promoting
osteogenesis.296
Tissue-engineered constructs for bone should be osteocon­
ductive to allow bone cells to adhere, proliferate, and migrate.
A range of ceramics (e.g., β-tricalcium phosphate and
hydroxyapatite) and polymers have been reported to meet
this requirement and a number have been used as bone sub­
stitutes clinically. Slowly degrading polymers, such as PLA
and PCL, have been widely investigated to produce tough
porous scaffolds for bone tissue engineering. FDA-approved
PCL-based scaffolds have been developed for bone tissue
engineering (Osteopore International), and used clinically as
burr plugs and sheets for orbital floor reconstruction in more
than 1000 patients.95,297 Altering the surface topography of
some of these more traditional bone substitutes, such as the
coating of ceramic titanium dioxide scaffolds with nanopar­
ticles made of the same material, may further augment their
regenerative potential.298

258
CHAPTER 16 • Tissue engineering
to maximize tissue assembly, and greater emphasis is now
placed on the critical role of vascularization in supporting
engineered tissue. With this integrated approach, there has
been an exponential growth in tissue engineering studies in
recent years, with a number of high profile clinical reports.315,316
Using our models of intrinsic vascularization, we have
been successful in generating a range of tissues including
cardiac tissue from rat neonatal cardiomyocytes,162 human
ASCs,317 and human iPSCs.318 The chamber has also been used
for the implantation of pancreatic islets,319 and liver progeni­
tor cells.165 We have also engineered tissue capable of secreting
growth hormone after implantation of mouse pituitary stem
shown in the idealized schematic in Fig. 16.29. The 3D arrange­
ment of chondrocytes is important to their function and the
ECM synthesis required for tissue development, and this may
be encouraged by suitable biomaterial scaffold design.69,308
Native articular cartilage is not isotropic but contains zones
of different ECM and cell organization and function. Current
research aims to develop biomaterial scaffolds which mimic
this feature of cartilage more effectively to achieve engineered
tissues with the normal zonal organization.305,309
A range of biodegradable polymers, including PLGA,
PCL, hyaluronan, gelatin, and collagen, have been used to
produce scaffolds and microspheres to support chondro­
genesis.82,306,308,310 Matrix-induced autologous chondrocyte
implantation is similar in procedure to routine autologous
chondrocyte implantation, but it relies on scaffold technology
where cells are expanded and cultured on porcine collagen
type I/III membranes prior to implantation.311 Hyalograft
(a form of the biomaterial product Hyaff-11 from FIDIA
Advanced Biopolymers, Italy) is a hyaluronic acid-based 3D
scaffold product for cartilage tissue engineering that has also
been used clinically.49,306
The main cartilage tissue engineering approaches are the
delivery of cells in a hydrogel or scaffold, the implantation of
acellular scaffolds tailored to stimulate chondrogenesis, or
scaffold-free techniques involving the culture and implanta­
tion of chondrogenic spheroids which may be derived from
stem cells such as iPSCs.312 The physical microenvironment of
the cells is important to tissue development, and can be
manipulated during in vitro culture using forces such as
gravity and hydrodynamic forces from the fluid motion to
affect the quality and size of cartilage tissue formed (Fig.
16.30).308,313 A number of bioreactor designs have been devised
to provide favorable in vitro growth conditions for chondro­
cytes.305,310 Other novel aspects being explored include 3D
printing, including the use of BioPen technology where cells
in an ECM carrier ink could be delivered through a pen
nozzle, allowing the filling of small surgical defects as well as
precise coverage over bone and cartilage surfaces.314
Tissue engineering of other organs
Striking advances are being made in identifying, character­
izing, and culturing various stem cells, in tailoring scaffolds
Fig. 16.29  A timeline for cartilage tissue engineering from in vivo implantation to
formation of the final tissue. (From Palsson BO, Bhatia SN. Tissue Engineering.
Upper Saddle River, NJ: Pearson Education, 2004.)
Cell signaling,
adhesion,
migration
Cell differentiation,
proliferation,
apoptosis
Matrix
remodeling
Implant
site
Cartilage
Bone
Second
Hour
Day
Week
Month
Year
Fig. 16.30  Effects of physical forces on tissue engineering: cartilage grown in vitro
in microgravity and in Earth’s gravity. (A) Tissue constructs obtained from growth in
rotating bioreactors on Earth for 3 months and in microgravity for 4 months. (B) On
Earth only for 7 months (tissues are stained with safranin-O to show
glycosaminoglycan in red). (C) Graph of tissue construct aggregate moduli and
glycosaminoglycan (GAG) content. (Based on data from Freed LE, Langer R, Martin
I, et al. Tissue engineering of cartilage in space. Proc Natl Acad Sci USA.
1997;94:13885–13890. © 1997 National Academy of Sciences USA.)
Sample
0
0.2
0.4
0.6
0.8
1
1.2
0
2
4
6
8
10
12
Aggregate modulus (MPa)
GAG content (% wet weight)
Natural
cartilage
Modulus
Earth
(3 months)
Earth
(7 months)
Earth
(3 months) +
microgravity
(4 months)
A
B
C

259
Case studies and status of tissue engineering for specific tissues
reseeding a decellularized tracheal graft with the patient’s
own epithelial cells and MSC-derived chondrocytes.330 A
5-year follow-up report noted that the graft remained patent,
vascularized, and completely recellularized with normal
function and no signs of rejection, although stenosis of the
native trachea near the anastomosis required repeated endo­
luminal stenting.331 This group has also developed a synthetic
tracheal graft fabricated from nanocomposite polymer, which
was reseeded with a patient’s cells in a bioreactor prior to
transplantation.332 Transplantation of bioengineered trachea
based on decellularized matrix has also been performed in a
paediatric patient.333 However, these remain individual case
studies and very much like with bladder tissue engineering,
expanded clinical studies with long-term follow-up is required
before these feats become an accepted treatment.
In 2014, a small pilot study of 4 patients with vaginal
aplasia involved the transplantation of vaginal organs created
from autologous epithelial and muscle cells obtained from a
vulval biopsy and seeded in segments of Surgisis (decellular­
ized porcine intestinal submucosa matrix).334 This technology
could be useful not only for vaginal reconstructions but also
for other soft tissue defects, although the lack of integrated
vascularization will likely limit the use to thin grafts implanted
into relatively well-vascularized sites.
Whilst clinical studies are still few and far between and
fraught with technical challenges, we are gradually moving
from thin flat tissues (skin, cornea) to tubular (trachea, blood
vessels) and hollow viscus organs (bladder), and in the fore­
seeable future will likely progress to solid organs such as the
kidney, heart or liver. Organ regeneration is the grail of tissue
engineering, and with the advent of iPS cells this field offers
a new direction in personalized organ repair – once the realm
of science fiction, but now yielding incredible discoveries at a
rapid pace.
cells into a vascularized chamber in mice320 and restored
thymus function in athymic rats after implantation of human
thymic tissue in vascularized chambers as evidenced by the
postoperative appearance of processed T cells in the blood
and by the rejection of skin allografts.321
Most recently, the field has placed particular interest in the
use of extracellular matrix scaffolds derived from decellular­
ized tissues and whole organs of various types, including
human heart,322 lungs,323 liver,324 and kidneys.325 Many pre-
clinical studies have demonstrated the recellularization of
these scaffolds with various cells including organ-specific
cells and other stem cells, with cell attachment and organiza­
tion into their native compartments to recapitulate tissue
architecture. Many of these bioengineered tissues and organs
have been implanted in small animal models, and several of
these areas have been translated into clinical studies.
An early human clinical trial on organ tissue engineering
was led by Atala et al. and focused on the bladder. A signifi­
cant milestone was the report in 2006 that engineered bladder
tissue showed promising outcomes in a clinical trial for cys­
toplasty reconstruction.326,327 Autologous cells from a bladder
biopsy were seeded on to biodegradable scaffolds and grown
in vitro prior to transplantation (Fig. 16.31). Bladder function
was shown to improve and this was sustained up to 5 years.
Atala provided an anecdotal note in 2014 that these patients
continued to do well up to 11 years later.328 However, a recent
Phase II trial by a separate group using the same protocol
reported no improvement in bladder compliance or capacity,
and serious adverse events such as bowel obstruction and/or
bladder rupture in 4 out of 10 patients.329 Atala considers the
disparity possibly to be due to a number of reasons. These
include variability in patient cells, batch-to-batch manufactur­
ing variability due to an individualized process, genetic
variability, and lastly difference in the clinical spectrum of
spina bifida patients in terms of their level of spinal cord
pathology, bladder innervation, and functional disturbance.328
While efforts are being made to address these unique issues
in regenerative medicine, it is worth remaining prudent and
cautiously optimistic while we await the results of the next
bladder trials before making any firm conclusions.
In 2008 Macchiarini et al reported the transplantation of the
first bioengineered trachea in a human patient, based on
Fig. 16.31  Tissue engineering of the bladder. (A) A biodegradable scaffold was seeded with autologous cells. (B) Anastomosed to the patient’s native bladder.
(C) Covered with fibrin glue and omentum. (From Atala A, Bauer SB, Soker S, et al. Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet.
2006;367:1241–1246. © Elsevier 2006.)
A
B
C
Bonus images for this chapter can be found online at

Fig. 16.3 (A) Thumb stump with distraction device. (B) X-ray of bone
distraction. (C) New bone filling the gap.

260
CHAPTER 16 • Tissue engineering
3. Cao Y, Vacanti J, Paige K, et al. Transplantation of chondrocytes
utilizing a polymer-cell construct to produce tissue-engineered
cartilage in the shape of a human ear. Plast Reconstr Surg.
1997;100:297–302. The paper reporting the spectacular image of the
“human ear” in the mouse which attracted much attention to the
potential of tissue engineering.
44. Engler A, Shamik S, Sweeney H, Discher D. Matrix elasticity
directs stem cell lineage specification. Cell. 2006;126:677–689.
52. Williams D. Essential Biomaterials Science. Cambridge: Cambridge
University Press; 2014. A recent textbook on biomaterials science which
addresses the field in detail, including considerations in the design,
selection, biocompatibility testing and use of biomaterials for tissue
engineering.
66. Go DP, Palmer JA, Mitchell G, et al. Porous PLGA microspheres
for dual delivery of biomolecules via the layer-by-layer assembly
technique. J Biomed Mater Res A. 2015;103:1849–1863.
90. Cosson S, Otte EA, Hezaveh H, Cooper-White JJ. Concise review:
tailoring bioengineered scaffolds for stem cell applications in
tissue engineering and regenerative medicine. Stem Cells Transl
Med. 2015;4:156–164. A recent review on scaffold design for tissue
engineering and stem cell culture with a focus on mimicry of the
biological microenvironment.
91. Cao Y, Mitchell G, Messina A, et al. The influence of architecture
on degradation and vascularisation of three dimensional
poly(lactic-co-glycolic) scaffolds in vitro and in vivo. Biomaterials.
2006;27:2854–2864. An in vivo study revealing the influence of scaffold
pore architecture on vascularization and tissue survival in vivo using the
commonly applied PLGA polymer as a biomaterial.
154. Tanaka Y, Tsutsumi A, Crowe DM, et al. Generation of an
autologous tissue (matrix) flap by combining an arteriovenous
shunt loop with artificial skin in rats: preliminary report. Br J Plast
Surg. 2000;53:51–57. This paper describes the first realization that the
“tissue” that grows into a space is not only vascular tissue and
demonstrates the tissue engineering concept where a vessel loop is
directed into a protected space that cannot collapse, such as a chamber in
a rat model.
156. Lokmic Z, Stillaert F, Morrison WA, et al. An arteriovenous loop in
a protected space generates a permanent, highly vascular,
tissue-engineered construct. FASEB J. 2007;21:511–522. This paper
provides a detailed analysis of angiogenesis from a vascular loop in a
dedicated space.
162. Morritt AN, Bortolotto SK, Dilley RJ, et al. Cardiac tissue
engineering in an in vivo vascularized chamber. Circulation.
2007;115:353–360. Captures the imagination and potential of tissue
engineering; significant because of the spectacular image of beating tissue
grown in a chamber in vivo.
163. Hussey AJ, Winardi M, Han XL, et al. Seeding of pancreatic islets
into prevascularized tissue engineering chambers. Tissue Eng Part
A. 2009;15:3823–3833.
174. Morrison WA, Marre D, Grinsell D, et al. Engineering of adipose
tissue in humans using the tissue-engineering chamber: a pilot
study. EBioMedicine. 2016.
180. Kolesky DB, Truby RL, Gladman AS, et al. 3D bioprinting of
vascularized, heterogeneous cell-laden tissue constructs. Adv
Mater. 2014;26:3124–3130. A demonstration of the potential to use 3D
printing to create vascularized tissue engineering constructs comprising
multiple cell types.
262. Jank BJ, Xiong L, Moser PT, et al. Engineered composite tissue as a
bioartificial limb graft. Biomaterials. 2015;61:246–256. A striking
proof-of-concept study on recellularizing decullarised tissue scaffolds
based on animal limbs using multiple cell types.
315. Lin X, Huang J, Shi Y, Liu W. Tissue engineering and regenerative
medicine in applied research: a year in review of 2014. Tissue Eng
Part B Rev. 2015;21:177–186. Outlines the recent applied research
activity in the tissue engineering field, with discussion of clinical trial
and significant advances.
Access the complete reference list online at

260.e1
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17 
Repair, grafting, and engineering
of cartilage
Wei Liu and Yilin Cao
SYNOPSIS
■ Cartilage is one of the most important tissue grafts in plastic surgery
and is widely used in auricular reconstruction, rhinoplasty, and facial
contouring.
■ This chapter introduces the background knowledge and surgical
skills required for harvesting cartilage graft and managing donor site
tissues.
■ Additionally, recent developments in engineered cartilage
fabrication and their potential application in plastic surgery are
described.
Introduction
Cartilage is a kind of connective tissue which is mainly com­
posed of chondrocytes and their extracellular matrices (ECM)
of type II collagen fibers, proteoglycans, and elastic fibers.
According to its composition, cartilage can be classified into
three types1: (1) hyaline cartilage; (2) fibrocartilage; and (3)
elastic cartilage. Fig. 17.1 shows the histology of the three
types of cartilages.
Hyaline cartilage is the most common type of cartilage and
can be found in costal, articular, tracheal, and nasal cartilage.
With the exception of articular cartilage, the free surface of
most hyaline cartilage is covered with perichondrium. In
addition to collagen II, hyaline cartilage is rich in glycosami­
noglycans and is thus characterized by its stiffness, that
permits sustained compressional loading.
Fibrocartilage is composed of bundles of thick collagen
fibers along with intervening unicellular islands of cartilage
arranged in small chains. Because of this unique structure,
fibrocartilage can provide high tensile strength and support­
ing function and is thus present in areas that are most subject
to frequent stress, such as meniscus, intervertebral discs,
symphyseal joints, and the joint portion of bone and
tendons/ligaments. In contrast to other cartilage, fibrocarti­
lage also contains type I collagen. In addition, it has moderate
amounts of proteoglycan but little glycosaminoglycans.
Elastic cartilage is characterized by its extremely high
elasticity because of the presence of abundant amounts of
elastic fibers. It is histologically similar to hyaline cartilage but
contains many elastic fibers which form an elastic fiber
network along with collagen fibers. This unique feature pro­
vides great flexibility so that elastic cartilage can withstand
repeated bending. It is mainly found in the outer ear structure,
and also in the larynx and epiglottis. It is also surrounded
with perichondrium.
The perichondrium is a layer of dense irregular connective
tissue that consists of two separate layers, an outer fibrous
layer and an inner chondrogenic layer. Collagen fibers and
fibroblasts constitute the fibrous layer. In contrast, the chon­
drogenic layer remains partially undifferentiated and is likely
to contain mesenchymal stem cells and chondrogenic pro­
genitor cells,2,3 which can play a role in cartilage repair and
regeneration.
Cartilage is a unique tissue with low metabolic rate due to
the sparsity of its cell population and its avascular structure.
The glycolytic activity and oxygen consumption of cartilage
approaches anaerobic condition and the tissue is nourished
by tissue fluid diffusion. Because of this unique feature, car­
tilage graft is relatively easy to survive when being implanted.
In terms of cartilage grafting, Sushruta Samhita in India prob­
ably is the first person to perform cartilage grafting in the
form of a composite graft.4,5 Cartilage graft is widely used for
repairing nasal or auricular defects as well as for reconstruct­
ing other tissues either as a pure cartilage graft or as a com­
posite graft. According to surgical procedures, cartilage can
be transferred either as a free graft or as a microvascular
composite graft. Generally, autologous cartilage graft will not
have metaplastic changes even being transferred to a different
location.6 In addition, grafting of free perichondrial tissue is
also a common procedure for cartilage reconstruction because

262
CHAPTER 17 • Repair, grafting, and engineering of cartilage
autologous cartilage grafting remains the most applicable
cartilage graft. Generally, common donor sites for harvesting
autologous cartilage grafts include auricular, nasal, and rib
cartilage.
Auricular cartilage graft
As an elastic cartilage, auricular cartilage is an ideal graft for
transplantation and perhaps is the most versatile of all carti­
lage grafts because it can be easily fashioned and contoured
into different shapes for various uses. Auricular cartilage can
be harvested easily under local anesthesia and a significant
portion of the concha can be removed without causing donor
site deformity.6
However, in practice, harvesting of the entire conchal car­
tilage is likely to cause collapsed conchal bowl, cymba concha,
and a horizontally short ear. To overcome this problem, Han
et al. proposed the surgical procedure that involves: (1) using
a postauricular incision to minimize visible scars; (2) harvest­
ing the entire cymba concha and cavum concha separately,
with at least 5 mm of the helical crus, leaving a lateral exten­
sion as a strut between them, as well as a 2-mm outer rim
along the conchal wall; and (3) by using a tie-over bolster
dressing that can serve as a mold for the conchal bowl. Fig.
17.2 presents the technique to avoid contour irregularity or
deformity after harvesting the maximal amount of conchal
cartilage graft.8
Auricular cartilage graft is often used as a framework
for ear reconstruction or auricular deformity correction, as
reported by Brent,9 Ono et al.,10 Firmin et al.,11 and others. In
addition, conchal cartilage can be used as a single-layered
graft for nasal, tarsal, and nipple reconstruction.6
The other important application of auricular cartilage is to
transfer as a composite chondrocutaneous graft for nasal
reconstruction. How to manage the donor site of auricular
composite graft aesthetically is important for its successful
application. Singh and Bartlett have summarized several
techniques to deal with the donor site of free auricular com­
posite grafts according to different applications12:
■ Technique 1: A composite graft that is less than 1 cm can
be harvested from the root of the helix and the dog ear is
run anteriorly toward the hairline with primary closure
of the defect (Fig. 17.3).
■ Technique 2: This is designed to close the donor site
wound of the composite graft with cartilaginous base of
1–1.5 cm, which is usually needed to repair a wider
defect of the alar rim (Fig. 17.4).
■ Technique 3: To repair a nasal defect with short vertical
dimension (height), the composite graft can be harvested
from the anterior aspect of the helical root and the dog
ear must be run superiorly and inferiorly. As shown in
Fig. 17.5, this technique can provide adequate tissue for
nasal reconstruction and an aesthetic closure without
disturbing the size of the ear.
■ Technique 4: For harvesting grafts with width of 1–1.2 cm
from the base of the helix, the defect may be closed
primarily by advancing the helical rim forward. The
possible resulting overprojection of the ear might be
addressed with a postauricular incision and a
scaphomastoid suture to match the prominence with the
contralateral ear (Fig. 17.6).
it has long been observed that perichondrium possesses the
ability to regenerate cartilage.7
Autologous cartilage grafts
and applications6
Although cartilage is considered as “immunologically privi­
leged”, and allogeneic cartilage may serve as a potential graft,
Fig. 17.1  Histology of (A) hyaline cartilage, (B) fibrocartilage, and (C) elastic
cartilage (silver staining). (Courtesy of Dr. Roger C. Wagner, Prof. Emeritus of
Biological Sciences at University of Delaware and Prof. Fred E. Hossler, East
Tennessee State University.)
A
B
C

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Autologous cartilage grafts and applications
Fig. 17.2  The surgical technique for harvesting the maximal amount of conchal cartilage graft. (A) Markings. It was necessary to leave at least 2 mm of the superior outer
rim along the conchal wall in order not to cause a noticeable change in the conchal bowl of the donor ear. Note the helical crus with its lateral extension between the
cymba concha and cavum concha. (B) The entire unit of the cymba concha and cavum concha was removed separately, leaving at least 5 mm of the crus helicis with an
extension to prevent collapsing. (C) There is almost no external evidence that the large conchal cartilage graft had been removed. (D) Two peanut cotton balls were placed
into the interstices of the ear and sutured in place with through-and-through 4-0 nylon sutures, which were removed 5 days later. (E) The harvested cymba and cavum
conchae were placed in situ to demonstrate their relationship with the intact helical crus and its lateral extension. (F) The surface area was 224.0 mm2 at the cymba concha
and 247.0 mm2 at the cavum concha. (Reproduced from Han K, Kim J, Son D, et al. How to harvest the maximal amount of conchal cartilage grafts. J Plast Reconstr Aesthet
Surg. 2008;61:1465–1471.)
A
B
C
D
E
F